Functional MRI at High Fields: Practice and Utility

Transcription

1 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 1 Functional MRI at High Fields: Practice and Utility Kamil Ugurbil, Wei Chen, Xiaoping Hu, Seong-Gi Kim, Xiao-Hung Zhu Center for Magnetic Resonance Research, University of Minnesota, MN, USA and Seiji Ogawa Bell Laboratories, Lucent Technologies, Murray Hill, NJ, USA 1 INTRODUCTION In the middle of the 19th century, a central debate about brain function revolved about the issue of whether the human brain processed all of its task as a single entity or function was compartmentalized. At about that time, an argument for the existence of regional specialization of human brain function was presented by Pierre Paul Broca. 1 Broca examined a patient who was unable to speak as a result of a stroke but was otherwise normal. Based on an autopsy performed subsequent to the patient's death, Broca concluded that the seat of the damage was an egg-sized lesion located in the inferior frontal gyrus of the frontal lobe in the left hemisphere; this general area is now commonly referred to as Broca's area, although its precise topographical extent remains ambiguous. This type of study and, later, intraoperative mapping efforts with electrodes were, until recently, the primary source of our information on functional compartmentation in the human brain. Recent techniques using NMR permitted the acquisition of such information much more rapidly and with greater spatial accuracy, fueling explosive developments in the investigation of human brain function. For example, the language area rst identi ed by Broca can now be visualized with unprecedented spatial resolution using functional MRI (fmri), in data collection times that last only a few minutes. Figure 1 displays the three-dimensional result of such a study. 2 The functional map generated is superimposed on the anatomical image during a covert language task. In these gures, the gray scale anatomic images are either opaque, permitting the visualization of activation only on the outer cortical surface, or rendered partially transparent so that activation in the interior of the brain and within its numerous folds (sulci) are visible, albeit with diminished intensity. These images are based on BOLD (blood oxygen leveldependent contrast), rst described by Ogawa in rat brain studies 3±5 and subsequently applied to generate functional images in human brain. 6±8 Today, images like those shown in Figure 1 can be generated by the 1.5 T scanners that are often found in hospitals, provided that the scanner is equipped with appropriate hardware to perform fast imaging. These are relatively `low' resolution images, in the several millimeter spatial domain, obtained with averaging over many executions of the same task by a single individual. However, there have been studies beyond this type of imaging using signi cantly higher magnetic elds. Shortly before the introduction of BOLD-based fmri, efforts were initiated in three laboratories, Universities of Minnesota and Alabama, and the National Institutes of Health, to explore the possibility of using high magnetic elds (4 T) for human MRI. However, these high- eld studies were initially met with skepticism, and the possibility of human imaging at magnetic elds much higher than 1.5 T was seriously questioned. This skepticism was not based on the existence of any experimental evidence; rather, it followed from concepts and theoretical considerations regarding the interaction of high-frequency electromagnetic waves with the conductive human body. These considerations had led prominent investigators in the eld of MRI research, as early as 1979, to suggest that human imaging would not be possible beyond 10 MHz (~0.24 T). 9 Of course, even clinical imaging is now performed at magnetic elds much higher than 0.24 T (up to 1.5 T), and recent work at 3 and 4 T has demonstrated that exquisite anatomical and functional imaging of the human head is achievable at these high elds. The efforts toward the development of fmri at our site was performed only at 4 T, coinciding with the effort to introduce and explore the use of these high magnetic elds. It is generally stated that the contrast-to-noise ratio (CNR) for the BOLD effect and hence for detection of activated areas increases at high elds, hence higher elds are better for fmri. This may indeed be the case for some eld strengths. A clear demonstration of this is illustrated in Figure 2, where mapping of the various functionally distinct regions within the visual cortex at 1.5 and 3 T depict larger areas of activation at the higher eld strength. These are images of the attened cortex. Similar functional maps have been published by Tootell and colleagues previously, mainly at 1.5 T. 10±15 However, the eld dependence of the BOLD-based fmri is rather complex. At high elds, for example, the fractional signal change coupled to alterations in neuronal activity may actually become smaller than that observed at lower magnetic elds such as 1.5 T; however, the speci city of signal changes detected by MR in relation to the actual site of neuronal activation may be substantially improved. Such trade-offs between CNR and speci city will also depend on the details of data acquisition. Therefore, consideration of the BOLD mechanism is imperative in understanding the underlying eld dependence of fmri. This chapter starts with such a discussion and subsequently gives a few selected examples of accomplishments that remain unique to high- eld fmri. 2 MECHANISTIC CONSIDERATIONS RELEVANT TO FIELD DEPENDENCE 2.1 Extravascular Blood Oxygen Level-dependent Effects Modeling of the effect of susceptibility gradients across vascular boundaries on MR signals have been considered by several groups, with similar conclusions. 16±24 If one considers an in nite cylinder as an approximation for a blood vessel with magnetic susceptibility difference, then the magnetic eld expressed in angular frequency at any point in space will be perturbed from the applied magnetic eld! Inside the For References see p. 18

2 2 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY Figure 1 Three-dimensional, whole brain images of activation during a language task based on word generation from a phoneme. Subjects were presented phonemes and were asked to think of as many English words as they could that contained the phoneme until the presentation of the next phoneme. The gray scale picture is the anatomical image. In color is the superimposed functional map. Brain images shown are (a) left hemispheres (b) right hemispheres, and (c) at an angle where the left hemisphere and top of the brain are both partially seen. The extensive frontal cortex activity depicted in (a) largely de nes the area of Broca and extends to area 46 as well. In (a)±(c), the anatomical images are opaque; consequently only activated regions that lie predominantly on the cortical surface are seen. Image (d) is identical to (c) except that the anatomical image was rendered partially transparent in order to `look through' and see the activated regions that would normally be blocked from the view by overlapping cortex (e.g., regions within sulci). In these particular views, activity in the medial part of the brain and other areas are also apparent, albeit with diminished intensity, in addition to the extensive activity in the left hemisphere frontal cortex. (With permission from Erhard et al. 2 ) cylinder, the perturbation,! B, will be given by:! in B ˆ Y! 0 fcos 2 1=3g 1 At any point outside the cylinder, the magnetic eld will vary, depending on the distance and orientation relative to the blood vessel and the external magnetic eld direction, according to the equation:! out B ˆ Y! 0 fr b =rg 2 sin 2 cos 2 2 In these equations, 0 is the maximum susceptibility difference expected in the presence of fully deoxygenated blood, Y is the fraction of oxygenated blood present, r b is the cylinder radius, and r is the distance from the point of interest to the center of the cylinder in the plane normal to the cylinder (Figure 3). Note that outside the cylinder the magnetic eld changes rapidly over a distance comparable to two or three times the cylinder radius; at a distance equal to the diameter of the cylinder from the cylinder center,! out B is already down to 25% of its value at the cylinder boundary. If such a blood vessel is present in a given voxel, the magnetic eld within this voxel will be inhomogeneous. The effect of this inhomogeneity in tissues can be understood in terms of dynamic and static averaging, the former arising as a result of the diffusive motion of the water molecules. First, let us ignore the blood in the intravascular space (i.e., inside the cylinder) and focus on the extravascular space only. In BOLD-based fmri, data are collected after excitation and an echo delay time TE in a gradient or a spin echo sequence. For list of General Abbreviations see end-papers

3 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 3 Figure 2 Mapping eccentricity and motion versus stationary stimulus in the visual cortex. Images are presented as a attened cortex. (a), (b) Data from an experimental mapping eccentricity; (c), (d) low-contrast moving stimulus versus stationary stimulus that activates areas MT and V3a. (a), (c) Data at 3 T, with statistical signi cances ranging from 10 5 to (b), (d) Data at 1.5 T analyzed with the same statistical criteria as in (a) and (c). Comparison of (a) and (c) and of (b) and (d) show the effects of magnetic eld strength at the same statistical signi cance. (Figure supplied by N. Hadjikhani and R. Tootell) Typical TE values used depend on the eld strength and the speci cs of the pulse sequence but, in general, range from ~30 to ~100 ms. If the typical diffusion distances during the delay TE are comparable to the distances spanned by the magnetic eld gradients, then during this delay the magnetic eld inhomogeneities will be dynamically time-averaged. Therefore, blood vessel size compared with the diffusion distances in this 30±100 ms time domain becomes a critical parameter in the BOLD effect (Figure 4). In this time scale, small blood vessels (e.g., capillaries) that contain deoxyhemoglobin will contribute to the dynamic averaging and result in a signal decay that will be characterized with a change in T 2 time. 17,18,25 In a spin echo experiment with a single refocusing pulse in the middle of the delay period, the dynamic averaging that has taken place during the rst half of the echo will not be recovered. Of course, applying many refocusing pulses, as in a Carr±Purcell pulse train, or applying a large B 1 eld (relative to the magnitude of the magnetic eld inhomogeneity) for spin-locking during this delay will reduce or even eliminate this signal loss through dynamic averaging. In a gradient echo measurement, For References see p. 18

4 4 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY Figure 3 A cylinder representing a blood vessel and the parameters that determine magnetic eld at a point outside of the cylinder when the susceptibilities inside and outside the cylinder are not the same dynamic averaging will occur during the entire delay TE. If the imaging voxel contains only small blood vessels at a density such that one half the average distance between them is comparable to diffusion distances (as is the case in the brain where capillaries are separated on the average by 25 m 26 ), then the entire signal from the voxel will be affected by dynamic averaging. In considering the movement of water molecules around blood vessels, we need not be concerned with the exchange that ultimately takes place between intra- and extravascular water across capillary walls. Typical lifetime of the water in capillaries exceeds 500 ms, 27±29 signi cantly longer than the typical T 2 and T2 values in the brain tissue and longer than the period TE typically employed in fmri studies. For larger blood vessels, complete dynamic averaging for the entire voxel will not be possible. Instead, there will be `local' or `partial' dynamic averaging over a subsection of the volume spanned by the magnetic eld gradients generated by the blood vessel. However, there will be signal loss from the voxel through static averaging if refocusing pulses are not used or asymmetric spin echoes are employed. Following the excitation and rotation onto the plane transverse to the external magnetic eld, a water molecule at a given point in space relative to the blood vessel will see a `locally' time-averaged magnetic eld that will vary with proximity to the large blood vessel. Therefore, the signal in the voxel will then be described by: S t ˆX c k e TE=T 2k e i! kte 3 k where the summation is performed over the parameter k, which designates small volume elements within the voxel, and c k is a constant; the time-averaged magnetic eld experienced within these small volume elements is! k in angular frequency units. The summation over k, therefore, covers the entire voxel. Because! k TE varies across the voxel, the signal will be Figure 4 Dynamic and static averaging regimes based on diffusion distances relative to the size of a compartment that differs in magnetic susceptibility from the surrounding tissue. The magnetic eld gradients are most prominent in the vicinity of the compartments with different susceptibility, i.e., red blood cells, capillaries, and large blood vessels. For large blood vessels, diffusion distances are not large compared with vessel radius, and hence they do not lead to dynamic averaging Figure 5 The susceptibility-induced R 2 (1/T 2 ) in the presence of water diffusion plotted as a function of cylinder radius at three different values of frequency shifts. TE is 40 ms and the fractional `blood volume' (i.e., volume within cylinder relative to voxel volume) is (With permission from Ogawa et al. 18 ) For list of General Abbreviations see end-papers

5 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 5 Figure 6 The fractional signal loss (S/S) for spin echo versus a gradient echo image acquisition as a function of cylinder radius TE, 40 ms; frequency shift owing to susceptibility difference, 40 Hz; fractional `blood volume' (i.e., volume within cylinder relative to voxel volume), (With permission from Ogawa et al. 18 ) `dephased' and diminish with increasing TE. This signal loss occurs from static averaging. In this domain, if the variation! k over the voxel is relatively large, signal decay can be approximated with a single exponential time constant T2. Figure 5 illustrates R 2 (i.e., 1/ T 2 ) as a function of the radius of the cylinder from modeling studies by Ogawa et al. 18 In this calculation, all other relaxation mechanisms that can contribute to transverse relaxation of water protons are ignored; only the effect of the susceptibility difference between the cylinders and their surrounding is considered. The imaging voxel was divided into many cubes (>10 6 ) with an edge dimension L, each of which contained a single cylinder of length L. The volume of the cylinder (i.e., `blood volume' b v ) relative to the volume of the cube is given by (r 2 b L)/L 3. Figure 5 displays the results as a function of cylinder radius and for a frequency shift ( 0 (1 Y)! 0 ) equal to 32, 48, and 64 Hz, b v =0.02 (simulating the capillary blood volume to total tissue volume ratio in the brain, 26 and TE of 40 ms (often used in studies at 4 T). Given the maximal susceptibility difference between fully oxygenated and fully deoxygenated blood 5 and taking Y to be 0.6, typical of venous blood in the brain, the frequency difference 0 (1 Y)! 0 is calculated to be 43 Hz at 4 T. The data in Figure 5 demonstrate that at a radius less than ~5 to 10 m, depending on the magnitude of the frequency shift, R 2 decreases because of dynamic averaging; the cylinder radius where this R 2 decrease becomes apparent is smaller at larger frequency shifts as expected. Above ~10 m, R 2 is approximately independent of the radius as the static regime dominates for all 0 (1 Y)! 0 values considered in these calculations. These modeling efforts also suggest that deoxyhemoglobincontaining microvessels in the brain [i.e., capillaries (5 m mean diameter) and venules (which can be 2±20 m in diameter 30 )] would be in the `dynamic' averaging regime. Note that the frequency difference between the intra- and extra-cylinder compartments depends both on the applied external magnetic eld B 0 and on the magnetic susceptibility difference between the cylinder and its surrounding. Consequently, either at very high magnetic elds or at very large values of, the radius at which a transition occurs from static to dynamic averaging will shift to values smaller than the capillary size. Such large values of can be achieved even at 1.5 T with bolus injections of contrast agents. Figure 6 illustrates the dynamic versus static averaging regimes and their dependence on cylinder radius (i.e., blood vessel radius) in a different way. In a spin echo study, static averaging will not come into play because the refocusing pulse will undo the `dephasing' and `rephase' the spins. Therefore, spin echoes will only be sensitive to small cylinder radii. For gradient recalled echoes, both dynamic and static averaging are, in principle, operative. However, for small cylinder radii, only the dynamic averaging regime will apply. Therefore, a plot of the ratio of fractional signal loss during TE (i.e. S/S) for a spin echo versus a gradient recalled echo will have a curve approaching unity as the cylinder radius approaches zero. At the other extreme, this ratio will diminish with increasing cylinder radius as dynamic averaging disappears and a static averaging regime dominates. The net result of these calculations (again so far not considering the intravascular effects) yields the following terms for contributions to R 2 : R 2 ˆ f 0! 0 1 Y gb vl large vessels 4 R 2 ˆ f 0! 0 1 Y g 2 b vs p small vessels 5 where and are constants,! 0 is the external magnetic eld in frequency units (rad/s) (i.e.,! 0 =B 0 ), 0! 0 (1 Y ) is the frequency shift owing to the susceptibility difference between the cylinder simulating the deoxyhemoglobin-containing blood vessel, b vl is the blood volume for large blood vessels (veins and venules with a radius greater than ~5 m for 4 T) and b vs is the small vessel blood volume (capillaries and small venules, less than ~5 m in radius, that permit dynamic averaging), and p is the fraction of active small vessels (i.e., lled with deoxyhemoglobin-containing red blood cells). An important feature of Equation (5) is the fact that it varies as the square of the external magnetic eld for small vessels where the effect is dominated by dynamic averaging. In contrast, the dependence on the external magnetic eld is linear for large blood vessels because they are in the static averaging domain. Note, however, that even for the capillaries the quadratic dependence of the extravascular BOLD effect on! 0 will not persist forever with increasing external magnetic eld. At some very high external magnetic eld strength, the frequency shift across the luminal boundaries of the blood vessel will be suf ciently large to displace even the capillaries into the static averaging domain, and hence to linear dependence on the applied magnetic eld. At the present time, there is no experimental evidence as to what that eld strength is, although results from our laboratory demonstrate that at 9.4 T there still exists an extravascular BOLD effect owing to dynamic averaging during activation (presented below). Another important point that can be surmised from Equations (4) and (5) is that the BOLD effect will be proportional to three physiologic parameters: regional cerebral For References see p. 18

6 6 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY blood ow (CBF) and oxygen consumption rate (CMRO 2 ) [since (1 Y)=CMRO 2 /CBF] and regional blood volume. Neuronal activity is coupled to all of these physiologic parameters. It has been suggested that regional CBF increases while CMRO 2 in the same area is not elevated commensurably, 31±33 resulting in decreased extraction fraction and lower deoxyhemoglobin content per unit volume of brain tissue. Another important prediction of the modeling studies is that there are large and small vessel extravascular BOLD effects. This has implications with respect to the speci city of functional images generated by the BOLD effect. While capillaries are uniformly distributed in tissue and are suf ciently high in density, large venous vessels are not; consequently BOLD effects associated with large vessels will not be as closely correlated with the actual site of neuronal activity. This issue will be discussed in greater detail below. 2.2 Intravascular Blood Oxygen Level-dependent Effects In the blood, hemoglobin is also compartmentalized within red blood cells and when the deoxy form is present, there are eld gradients around the red cells. However, because the dimensions of red cells are very small compared with diffusion distances, the effect is dynamically averaged and becomes a T 2 effect only. The dynamic averaging in this case also involves exchange across the red blood cell membrane, which is highly permeable to water. Consequently, in the presence of deoxyhemoglobin-containing red blood cells, the T 2 value of blood decreases. This effect was noted by Thulborn and was shown to increase quadratically with eld strength as expected from dynamic averaging owing to diffusion in the presence of eld gradients. 34,35 Therefore, even when we neglect the extravascular effect described above, the T 2 time of blood itself will change when the deoxyhemoglobin content is altered by elevated neuronal activity, and this will lead to a signal change in a T 2 -ort2 -weighted image. This effect will be present wherever the deoxyhemoglobin content has changed, potentially both in large and small blood vessels. As discussed previously, an increase in cerebral blood volume (CBV) results in signal loss in the extravascular BOLD effect. If the extravascular BOLD effect is neglected and only the intravascular contribution is considered, the increase in CBV with elevated neuronal activity can be more complex. Essentially, the ratio of blood to tissue volume will increase in a given voxel when CBV is elevated during increased neuronal activity. If the blood T 2 time is longer than that of tissue, then the CBV increase will lead to a signal increase rather than the decrease that is always expected from the extravascular BOLD effect. At 1.5 T, blood T 2 can be calculated from: 36 1=T 2 ˆ4:02 41:5 1 Y 2 which yields ~250 ms for arterial blood (Y&1), and 94, 129, and 176 ms for venous blood with Y=0.6, 0.7, and 0.8, respectively. Experimental determinations give human arterial and venous blood T 2 times of ms and ms, respectively, at 1.5 T. 37 Compared with these blood T 2 values, cerebral tissues display T 2 values that range from ~70 to 90 ms at 1.5 T, 38 where the longest value is associated with cortical gray matter. These investigators measured T 2 values at 1.4 T. These values are expected to be also valid for 1.5 T given the 6 relatively slow variation of T 2 with magnetic eld strength for water in the brain. For example T 2 of cortical gray matter is ms at 4 T 39 as opposed to 87 2 ms reported for 1.4 T. 38 At 1.5 T, the value of T 2 of both arterial and venous blood is longer than that of gray matter where the blood volume and alterations in blood volume coupled to neuronal activity are most signi cant. Taking into account that the coef cient 41.5 in Equation (6) should be proportional to the square of the static eld (except at very high magnetic elds), the T 2 time for venous blood can be calculated at other eld strengths. At the same Y values, the predicted blood values for T 2 are approximately 4, 7, and 14 ms, respectively, for 9.4 T, in agreement with the experimentally measured values of ~7 to 8 ms in rat venous blood for a Y value in the 0.7 to 0.8 range (unpublished data from our laboratory). This rapid decrease in venous blood T 2 with increasing magnetic elds has signi cant implications for high- eld fmri studies and will be discussed below. At 4 T, where high- eld human fmri studies have so far been performed, blood T 2 is estimated to be 20, 31 and 63 ms for Y=0.6, 0.7 and 0.8, respectively, using Equation (6); these values are comparable to or signi cantly less than the gray matter T 2 time of ms. 39 Blood contribution comes into the BOLD phenomenon in a second special way when blood occupies a large fraction of the volume of the voxel, in other words when a large blood vessel occurs in the voxel. When deoxyhemoglobin is present in the blood, the blood water will result in dynamic averaging in the gradients surrounding the red blood cells and it will behave as if it encounters the uniform magnetic eld given by Equation (1). This will differ from the magnetic eld experienced by the rest of the voxel. In the immediate vicinity of the blood vessel, the magnetic eld will vary and it will approach a constant value in tissue distant from the blood vessel. For simplicity, we can neglect the gradients near the blood vessel and consider the voxel to be composed of two large bulk magnetic moments, one associated with blood and the rest with the extravascular volume. These magnetic moments will precess at slightly different frequencies, the difference in frequency given by Equation (1); therefore, the signal from the voxel will decrease with time as the two moments lose phase coherence. In this scenario, the signal can even oscillate as the phase between the two magnetic moments increases and then decreases. We will refer to this as a type 2 blood effect in fmri. When a voxel only contains capillaries, the blood volume is ~2%; 26 hence, the type 2 blood effect cannot exist for such a voxel. However, when a large blood vessel or vessels are present in the voxel, blood volume can signi cantly increase and become comparable to or even exceed the tissue volume in the voxel. If, of course, the voxel is smaller than the blood vessel dimensions, and the entire voxel is occupied by blood, then here too the type 2 effect does not come into play. Note that the type 2 blood effect and the extravascular BOLD effect in the static averaging regime are similar in nature because they both involve signal modulation owing to dephasing of magnetization within a voxel. The main difference is the presence of a large blood component and the lack of suf cient variation in resonance frequency within the voxel to approximate the signal modulation as an exponential signal decay in the type 2 blood effect. Similar to the extravascular BOLD effect in the static averaging domain, the type 2 blood For list of General Abbreviations see end-papers

7 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 7 effect will be refocused in a symmetric spin echo and thus nulli ed. Consequently, it will not be present in purely T 2 - weighted BOLD and functional images derived from it. Another important point is that this type 2 effect will diminish and even disappear at high magnetic elds because the T 2 value of blood gets very short and blood signal contribution becomes negligible. 2.3 Blood Oxygen Level Effects Detected in T 2 and T 2 by fmri The origin of the signal intensity changes that are detected in T 2 versus T2 -based BOLD fmri images, and hence the dependence on magnetic eld of these effects, differ signi cantly, as can be surmised from the above discussion. These differences are summarized here. The T2 -based BOLD signal can arise from both intravascular (blood) and extravascular effects originating from both large and small blood vessels. The relative contributions of these effects will depend on the magnetic eld strength. In a T 2 - based BOLD fmri map, the signal changes come from intravascular effects (i.e., blood T 2 changes), hence from both large and small blood vessels, and from an extravascular effect associated only with microvessels such as capillaries and small venules. The major difference is that the extravascular BOLD effect in a T 2 image can only arise from the microvasculature, whereas in T2 images, it can originate from blood vessels of all sizes. It is often suggested that T 2 -based fmri avoids a large vessel contribution. This claim is not correct because it ignores the intravascular blood-associated BOLD effect. At eld strengths of 1.5 T, or even 4 T, there clearly exists a signi cant blood contribution to the BOLD effect during increased neuronal activity. At 1.5 T, T 2 -based fmri maps are predominantly if not exclusively af liated with blood T 2 changes, hence with intravascular space, as discussed in the next section. Such blood contribution can originate from both large and small blood vessels. Only at very high elds (e.g., 9.4 Tesla) is T 2 - based BOLD fmri largely associated with the capillaries because intravascular blood contributions are suppressed by the very short T 2 value of blood (discussed in greater detail in the next section). 2.4 Experimental Studies Early in the history of fmri, it was demonstrated experimentally that the BOLD effect in human brain functional maps contained contributions from macroscopic venous blood vessels (i.e., vessels of ~0.5 mm or larger that can be visualized in MRI images) as well as from regions where no such macroscopic vessels were identi ed. 40 This has profound implications for the effective spatial speci city and spatial resolution that can be achieved with fmri since large blood vessels do not exist at high density in the brain, where capillaries are common. Therefore, functional maps based on the macrovasculature can be signi cantly distant from the actual site of increased neuronal activity and thus misleading if high-resolution mapping in the millimeter or smaller scale is desired. The effect of the blood vessel size on BOLD will depend on magnetic eld strength and the relative contributions of intra- and extravascular effects. These relationships can be experimentally evaluated. The issue of extra- and intravascular BOLD effects has been experimentally examined using a Stejskal and Tanner 41 gradient pair of either the same or opposite polarity, depending on whether gradient or spin echoes are used, respectively. Such gradients were rst used to examine molecular diffusion; therefore, their use to alter the image signal intensity is often referred to as `diffusion weighting' even though there are additional perturbations that arise from the use of such gradients. In experiments employing the Stejskal±Tanner gradients, the important parameters are the magnitude and the duration of the gradient pulses and the time separation between them. Frequently, the results are evaluated in terms of a parameter b, which is equal to (G) 2 ( /3) where is the gyromagnetic ratio (rad s 1 G 1 ), G is the magnetic eld gradient magnitude (G cm 1 ), is the duration of the gradient pulse, and is the separation in time of the onset of the two gradient pulses (where 1 G=10 4 mt). In simple isotropic diffusion, the MR signal in the presence of Stejskal±Tanner gradient, decays according to exp( bd) where D is the diffusion constant. For owing spins, b does not have such an immediately obvious physical meaning. If there is ow, the spins will acquire a phase that will depend on their velocity along the direction of the gradient and the gradient magnitude. Within vasculature, however, blood velocities are not uniform especially for vessels of large diameter. Furthermore, the blood vessels may change directions within a voxel, and there may be several different blood vessels with different ow rates and/or different orientations relative to the gradient direction. Since the blood signal detected from the voxel will be the sum of all these, the net result can be nulling of signal from owing spins through dephasing. The ability to nullify the intravascular signals by use of Stejskal±Tanner pulsed gradients provides the means to distinguish between intra- and extravascular BOLD effects in functional images. Such experiments have been performed at 1.5 and 4 T on humans. The studies at 1.5 T have indicated that most of the BOLD-based signal increase during elevated neuronal activity is eliminated by Stejskal±Tanner gradients, leading to the conclusion that most of the fmri signal at 1.5 T arises from intravascular effects. 42,43 This intravascular BOLD effect may be associated with macroscopic blood vessels since it is debatable whether the gradient pulses used can suppress intravascular signals from microscopic blood vessels such as capillaries and small venules. 44 As discussed above, the intravascular BOLD effect in functional brain imaging can arise from the change in blood T 2 time or from what we described as the type 2 blood effect. The latter is refocused by symmetric spin echoes and, therefore, can only be present in gradient recalled echo or asymmetric spin echo measurements. Given the fact that symmetric spin echo fmri experiments at 1.5 T yield very weak effects compared with gradient recalled echo studies, the data acquired with suppressing the blood component suggests that most of the fmri signal at 1.5 T arises from type 2 blood effects. The same conclusion was reached in high-resolution two-, or three-dimensional gradient recalled echo studies of motor cortex activation. 45 The effect of the Stejskal±Tanner gradients on brain tissue signal intensity at 4 T is illustrated in Figure 7a for a spin echo sequence. When activation studies were performed with such gradients at this eld strength, ~60% of the `activated' pixels disappeared at small b values but the remaining pixels persisted For References see p. 18

8 8 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY Figure 7 (a) The effect on signal intensity of variations in the value of b in the human brain under diffusion weighting in a spin echo sequence. (b) Number of activated pixels in the human visual cortex during visual stimulation at 4 T before and during application of bipolar gradients with increasing b value. The sequence was a spin echo sequence yielding T 2 -weighted BOLD images. Residual signal intensity that persists at high b values can come only from the microvasculature. Data were from four subjects (*) or one subject (*). (With permission from Menon et al. 46 ) as the gradient strength was increased to very large b values (Figure 7b). This suggests that at 4 T extravascular and/or capillary-level intravascular BOLD effects exist during activation as well as a signi cant intravascular contribution associated with the macrovasculature. Note that we have grouped extravascular and capillary-level intravascular effects together because it is not clear whether these gradients are suf- cient to nullify the intravascular signal from capillaries. At 9.4 T, the effects of the Stejskal±Tanner gradients become even more interesting. In a T 2 -weighted fmri study conducted in the rat brain (forepaw stimulation, symmetric spin echo with one 180 pulse), activation does not alter at all with changes from very small to very high b values (Figure 8). 47 The T 2 -based BOLD effects can only come from the blood through a change in the blood T 2 or from extravascular effects associated with capillaries and comparably sized venules. The gradient pair will suppress the blood effect, except possibly in capillaries and postcapillary small venules. Therefore, it can be concluded that at this very high magnetic eld there exists a strong and dominant BOLD effect originating from microscopic vessels. Intravascular effects associated with a blood T 2 change is a priori not expected to be signi cant at this eld strength because the T 2 value of venous blood is very short at 9.4 T (~5 ms), 47 as discussed above. Even arterial blood has a short T 2 time at this high magnetic eld strength (~30 ms) FUNCTIONAL IMAGING BASED ON CEREBRAL BLOOD FLOW BOLD contrast relies on the interplay between CBF and CMRO 2 as well as blood volume, and, as such, it represents a complex response controlled by several parameters. 18,21,23±25,48,49 Recent MR techniques, however, can also generate images based on quantitative measures of changes in CBF coupled with neuronal activity. 7,50±54 These CBF techniques rely on tagging the blood spins differentially within and outside of a well-de ned volume. For example, in the FAIR technique 51±53 frequency-selective inversion pulses are used to invert the longitudinal magnetization within a `slab' along one direction (typically axial); in the absence of blood ow, the spins relax back to thermal equilibrium only by spin-lattice relaxation mechanisms characterized with the time constant T 1. If ow is present, however, the relaxation becomes effectively faster as unperturbed spins outside the inverted slab ow in and replenish the net magnetization within the slab. Consequently, the effective spin-lattice relaxation in FAIR as well as other, similar ow-sensitive techniques 48,51,53±58 becomes characterized by a shorter time constant, T1, which is related to blood ow. It also follows naturally that if the inversion pulse in FAIR does not de ne a slab but inverts everything in the whole body (i.e., it is nonselective), blood ow does not enter into the problem. Therefore, in the FAIR technique, two images are acquired consecutively, each after a xed delay period subsequent to the inversion pulse; in one, the inversion pulse is slab selective and in the other it is nonselective. The difference image generated from this pair is a ow-sensitive image. One of the unique aspects of CBF-based functional maps is that macrovascular ow components can be selectively suppressed while sensitivity to microvascular ow and tissue perfusion changes is enhanced. In an experimental approach such as FAIR, this selective microvascular sensitivity is accomplished by changing the delay time after the initial inversion pulse and the subsequent signal excitation before image acquisition. In CBF-based functional imaging, longer delays (>1 s) emphasize microvascular ow and perfusion whereas shorter delays yield predominantly large vessel images. 51,59 The former, of course, is highly desirable for generating functional maps. Herein lies the magnetic eld dependence of CBF-based functional images. The T 1 time of tissue water gets signi cantly longer at high magnetic elds. Consequently, it is easier For list of General Abbreviations see end-papers

9 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 9 Figure 8 Diffusion-weighted spin echo fmri maps at 9.4 T with b values of 6.1 (a) and 438 s mm 2 (b) overlaid on one of the original consecutively acquired echo planar images (BOLD and diffusion weighted) collected during the functional imaging study. Coronal single-slice singleshot spin echo planar images of rat brain were acquired with a matrix size of 64 32, a Field of View of 3.0 cm 1.5 cm, a slice thickness of 2 mm, and TE of 30 ms. Somatosensory stimulation was used. The color bar indicates a maximum cross-correlation value from 0.7 to 0.9. Signal intensity (shown in background) was signi cantly reduced by bipolar gradients, as expected owing to diffusion. Localized activation is observed at the somatosensory cortex in the contralateral side of a stimulated forepaw. Foci of activation site (color) agree very well in both fmri maps. (c) A Turbo FLASH image shows the region of interest. (d) Time courses of diffusion-weighted images within the region of interest. If the macrovascular contribution was signi cant, relative BOLD signal changes would decrease when a higher b value was used. However, relative signal changes remained the same in both images, suggesting that extravascular and microvascular components predominantly contribute to spin echo BOLD at 9.4 T to detect the `microvascular' ow and perfusion and suppress the macrovascular component (see Tsekos et al. 59 ). In view of the macrovascular problems that can deleteriously affect BOLD images, the question arises as to why CBF-based images are not the preferred approach in generating functional maps. The answer is that the CNR is higher in BOLD images, and, in particular, rapid acquisition techniques covering the whole or a large subsection of the brain are yet to be developed with CBF-based methods. Unlike the CBF-based techniques, however, BOLD images can be acquired rapidly over the whole brain. Consequently, this approach remains the main method employed in fmri applications. 4 SPATIAL SPECIFITY In fmri studies, there are two reasons for concern regarding spatial speci city; one is the sensitivity to different size blood vessels and the presence of a macrovascular contribution, and the second is the spatial speci city of the physiologic and metabolic events that ultimately yield the functional images. 4.1 Early Blood Oxygen Level-dependent Responses Optical measurements in animal experiments have demonstrated that the onset of task-related activation rst results in signal changes interpreted as caused by an increase in deoxyhemoglobin content. 60,61 This deoxyhemoglobin increase peaks at ~3 s after task onset and is subsequently reversed, ultimately resulting in a relatively large decrease in overall deoxyhemoglobin content. If CMRO 2 is elevated because of energy requirements of increased neuronal transmission, oxygen extraction and consequently the deoxyhemoglobin level will also be elevated provided the blood ow remains constant. In contrast, a CBF increase alone without any alterations in CMRO 2 will cause only a decrease in deoxyhemoglobin; if there is a difference in the metabolic and hemodynamic response times of these two processes, with the latter lagging behind, the time dependence of deoxyhemoglobin content in the `activated' region will resemble the biphasic curve illustrated in Figure 9. Presence of early deoxygenation has been recently reported based on direct measurements of blood partial pressure of oxygen, 62 indicating that there indeed exists an For References see p. 18

10 10 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY Figure 9 Hypothetical time courses of cerebral blood ow (CBF) and oxygen consumption rate (CMRO 2 ) that would predict an early negative BOLD response early increase in CMRO 2 that precedes the onset of enhancement in blood ow. The early deoxyhemoglobin increase can also arise from a rapid elevation in blood volume before the onset of blood ow augmentation; such a volume increase would actually result in elevated oxy- and deoxyhemoglobin contents. A subsequent rush of oxygenated blood through enhanced blood ow would then decrease the deoxyhemoglobin level, provided oxygen utilization does not increase commensurately with CBF. The possible occurrence of such a rapid volume increase has been suggested by recent studies in the rat brain. 63 Malonek and Grinvald have argued that the early response is spatially more speci c and de nes better the columnar structure in the cat visual cortex examined in that study. 61 In contrast, they suggest that the CBF increase is not as speci c, ooding not only the active but also inactive columns and surpassing in spatial extent the actual area of activation by several millimeters. These claims suggest that the BOLD effect associated with the hyperoxygenated, high- ow phase during increased neuronal activity will also be spatially nonspeci c over several millimeters, re ecting the spatial distribution of the CBF response. The question arises whether the early response associated with increased deoxyhemoglobin can also be detected as a negative signal change in BOLD-weighted MRI and whether this response can be used as a means of obtaining functional images with the MR approach. An early negative response to activation was rst reported by Hennig and colleagues using a MR spectroscopy method, monitoring signal from a relatively large voxel but with high temporal resolution. 64 Subsequently, rst using averaging of data from a number of subjects 65 and later with single subjects, imaging studies documented the presence of a small but detectable early negative signal change, a `dip', at high magnetic elds. 66±70 Recently, the presence of this initial response was also observed in fmri studies conducted in monkeys at 4.7 T; 71 the magnitude and the time course was very similar to the human data obtained at 4 T. Figure 10 illustrates the time dependence of BOLD contrast fmri signal changes observed in the human visual cortex during and subsequent to a brief visual stimulation that lasted 2.4, 3.6, and 4.8 s. 67 In this study, T2 -weighted, gradientrecalled echo planar images were acquired rapidly covering only a few slices in the visual cortex, thus sacri cing spatial extent of coverage and spatial resolution in favor of time resolution. Furthermore, the brief stimuli were repeated several times (four to ten) and images were collected in synchrony with the stimulus presentation so that they could be averaged. The data revealed that during and following the brief visual stimulation period the BOLD-based fmri signal initially decreased; this decrease was reversed at about 3 s, resulting subsequently in a large signal intensity increase (Figure 10). For the longer stimulation period, a signal intensity decrease below baseline was also seen towards the end of the time course; this `late-phase', poststimulation decrease has been observed before, even in the very rst fmri papers, and may re ect a difference in the post-task temporal responses of blood volume 72 and/or CMRO 2 values 73 returning to basal levels much more slowly than did the blood ow. Figure 11 displays visual stimulation activation maps in the sagittal plane constructed from these data with the negative signal changes color coded in blue/purple colors and the positive response in red/yellow colors. Examining the images constructed from data collected during the early (negative) and the late (positive) response, we see that the early response is restricted to the anatomically well-de ned visual area V1 (primary visual cortex) along the calcarine ssure while the later `positive' BOLD image displays apparent `activation' in areas distant from this region; in this study, these distant areas represent artifacts associated with the MR methodology, namely the macrovascular in ow effect, which has been already discussed. The macrovascular in ow effect was not suppressed in these images because they were rapidly acquired and were obtained using a surface coil both for signal detection and excitation; consequently, it was practically impossible to achieve the condition of `full relaxation' between images that would have eliminated these artifacts without signi cantly sacri cing signal-to-noise ratio (SNR). The macrovascular in ow effect does not appear in the early images because it is absent owing to the slower hemodynamic response time. If the macrovascular effects are ignored and the functional maps around the calcarine ssure are compared, they are similar for the early negative and the later positive BOLD images. This observation has implications with respect to the spatial speci city of the functional maps generated by fmri. It suggests that, provided the macrovascular `in ow' artifacts are eliminated, the spatial speci city of the early negative BOLD-based maps and the later CBF-dominated positive BOLD images are very similar at the resolution of this study (~3 to 4 mm isotropic). However, this study does not yet answer the question of spatial speci- city raised by Malonek and Grinvald 61 because the study is not conducted at a suf ciently high spatial resolution. It is likely that higher magnetic elds will be necessary to achieve greater spatial resolution with this early response because it is a very weak effect. The fact that it is detectable by BOLDbased fmri, however, is signi cant both from a mechanistic point of view and for future developments in fmri. The early negative response has been examined further 74 and demonstrated to be linearly dependent on TE. The percentage signal change varied from individual to individual in the nine subjects studied, ranging between 0.38 and 0.95% at 21 ms TE and 0.93 and 3.2% at 45 ms TE at 4 T; however, in For list of General Abbreviations see end-papers

11 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 11 Figure 10 Signal intensity time course for activated areas in the primary visual cortex (V1) for three different periods of brief visual stimulation. Initially, the signal decreased to a valley, which is the early negative response or dip; subsequently, the signal increased leading to a peak that corresponded to the positive BOLD effect employed in functional images. This positive BOLD effect is much larger in magnitude than the early negative BOLD effect. Note that the poststimulation undershoot is seen prominently only when the visual stimulation duration is 3.6 ms or larger. (With permission from Hu et al. 67 ) each case, a linear dependence on TE was strongly evident. Figure 12 illustrates this linear dependence in a way that accounts for the intersubject variability. The presence of this linear response with TE is a priori expected if the initial negative change arises from a BOLD effect re ecting an increase in regional cerebral deoxyhemoglobin content. If the early response arises from an elevated CMRO 2 prior to the onset of an increase in blood ow, then initial stages of this early response will be associated with the capillary bed. At later times, this deoxyhemoglobin alteration will show up in the venules and veins as a result of blood ow. Capillary BOLD effects scale as the square of the magnetic eld, which would explain why this early negative response has not been detected or has been reported to be very small at 1.5 T. A recent study in our group was able to identify a small early response at 1.5 T in the visual cortex; comparison with the 4 T data suggests that the ratio of the peaks corresponding to the early negative and subsequent positive responses increased linearly with eld strength. Given the large macrovascular contribution at both 1.5 and 4 T to the late hyperoxygenated state, only a linear dependence on B 0 is expected for this positive BOLD response. This suggests that the early response must increase quadratically with the magnetic eld. 4.2 High-Resolution Imaging The issue of spatial speci city and resolution of fmri can also be addressed using speci c experiments to map functionally distinct structures with well-de ned organization and topography in the human brain. Early experiments introducing the fmri methodology employed such a strategy and examined the hemispheric lateralization in brain function. 6±8,75 For example, simple motor tasks are expected to be lateralized ipsilaterally in the cerebellum; 75 in other words, a simple motor task with the left or the right hand should predominantly activate the left or the right cerebellar hemisphere, respectively. This is in direct contrast to what is expected and observed in V1. However, detection of functional specialization with respect to hemispheric laterality only reveals the existence of a very coarse level of spatial speci city in a spatial domain that can be characterized as several centimeters; certainly, this level of speci city was demonstrated in the very rst studies introducing fmri. 6±8 On a much ner spatial scale (e.g., millimeter and submillimeter), it is possible to examine activation of small subcortical nuclei and even that of lower-level neuronal functional organizations such as the ocular dominance columns (ODCs). For References see p. 18

12 12 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY Figure 11 Functional images of visual stimulation constructed from (a) the early negative response (color coded in blue) and (b) the peak of the subsequent positive response (color coded in yellow/red). (With permission from Hu et al. 67 ) The thalamus provides an excellent case for evaluating the question whether structures that are only a few millimeters in size can be accurately mapped by fmri methodology. The thalamus contains several distinct, anatomically well-de ned regions or nuclei. These nuclei serve as relay points for a remarkably large number of pathways. For example, retinal output projects mainly to the lateral geniculate nucleus (LGN), a small, subcentimeter nucleus located posteriorly and ventrally within the thalamus. In turn, the LGN activates V1. 76±80 Detection of LGN activation by fmri has previously been reported at 1.5 and 4 T eld strength. 81±83 The LGN is functionally and spatially segregated and compartmentalized. Each of six LGN layers receives inputs from the speci c visual eld via the retina and then retinotopically signals to V1. 84 The speci city and resolution of fmri can be examined by testing the feasibility of mapping the retinotopic organization within this small nucleus. We have recently conducted such a study using a checkerboard visual stimulus covering different sections of the visual eld. 85 In particular, either upper or lower hemi eld stimulation against a dark control period was utilized. This is expected to activate LGN bilaterally but not uniformly. Spatial differentiation re ecting the retinotopical relationship between the upper and lower visual elds was detected and distinguished within the LGN. This is illustrated as composite maps for four individual subjects in Figure 13. The LGN activation induced by the upper visual eld stimulation (green and red pixels) was more inferior in location (closer to the hippocampal formation) compared with that induced by the lower visual eld stimulation (yellow Figure 12 Echo time dependence of the early negative response (dip) to activation. For each of nine individuals studied, an average percentage change was calculated for the three echo delays used. Subsequently, the percentage change for the dip at each echo time was divided by this average for that individual. Then, the `normalized' percentage change at each echo time was averaged across the nine subjects studied. The data points represent the mean and standard deviation obtained from this procedure. The line is a linear t to the data. (With permission from Yacoub et al. 74 ) For list of General Abbreviations see end-papers

13 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 13 Figure 13 High-resolution fmri mapping of lateral geniculate nucleus activation during the upper- eld and lower- eld red/black checkerboard visual stimulations (illustrated at the top of the gure). Images are in the coronal orientation. The activations in color represent composite fmri maps obtained by combining the activation maps generated by upper-visual eld versus a dark control period and lower visual stimulation versus a dark control period. The red pixels represent the overlap between these two activation maps. Yellow and green identify pixels activated only by the lower- and upper- eld stimulation, respectively. (Adapted from Chen et al. 85 ) and red pixels). This relationship is consistent with data obtained in the non-human primate visual system, which has been extensively studied using microelectrode recording 86 and selective lesions. 87,88 Our results also illustrate that the LGN in humans behaves similarly to V1. 10 However, the upper and lower visual eld representations in V1 are anatomically separated by the calcarine ssure and are distinguishable without overlap. In contrast, they are continuous in LGN layers. This contributes to the partial overlap of LGN activation between the upper and lower visual elds (Figure 13). From the LGN, geniculostriate projections to V1 continue to carry left or right eye input separately and terminate in layer IVC of V1, where they are arranged in a system of roughly parallel alternating stripes known as ODCs. In non-human primates and other vertebrates (e.g., cats), the organization of these columns has been studied by histological stains, autoradiography, and microelectrode recordings 80,89,90 and by optical imaging of intrinsic signals. 60,61,91±93 In humans, the ODCs have been demonstrated at autopsy in the striate cortex by histochemical staining for cytochrome oxidase; 94,95 however, a noninvasive technique for examining human striate cortex organization on the scale of cortical functional subunits has not been available. The hemodynamic-response mechanism that allows visualization of orientation columns and ODCs in awake monkeys by optical imaging of intrinsic signals demonstrates that corticovascular responses to visual stimuli can be localized to the columnar level in several mammalian species. In particular, the optical data demonstrate that, while the CBF response may not be speci c at the ODC level, 61 a deoxyhemoglobin difference across the active and inactive columns is generated presumably because of the enhanced CMRO 2 in the active but not the inactive column (Figure 14). This deoxyhemoglobin difference between the two columns should, in principle, be detectable by BOLD-based fmri provided the technique has suf cient speci city as well as sensitivity (SNR) to achieve the required high spatial resolution. For example, if large vessel contributions dominate the BOLD contrast observed, speci city will be inadequate to detect ODCs. For fmri studies using clinically available hardware, ~3 mm in-plane resolution and ~5 mm slices are typical because of the limited SNR available without extensive data averaging. Using the same high-resolution fmri pulse sequence with imaging hardware and parameters optimized at three different eld strengths, we have found that the SNR at 4 T is at least four times higher than at the much more commonly available 1.5 T For References see p. 18

14 14 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY Figure 14 Deoxy- and oxyhemoglobin response as a function of time during a brief period of visual stimulation in active and inactive columns in the cat visual cortex. The gure represents a summary of ndings by Maloneck and Grinvald. 61 eld strength. 96,97 This increase is suf ciently large to attempt imaging of ODCs in human V1, which are approximately 0.8± 1 mm on a side for a column and 5±10 mm long. 94,95 Using a simple visual paradigm in combination with an optimized rf coil, head restraints, subvoxel image registration, and the enhanced SNR provided by 4 T eld strength, it has been possible to demonstrate adjacent image pixels in human V1 that respond primarily to left or right eye photic input. 98±101 Figure 15a demonstrates a magni ed picture of cortical ribbon along a sulcus in V1. The image plane is along the calcarine ssure and columns appear as rectangles of approximately 1 mm1 mm separated by ~1 mm in the cortical gray matter. The color map represents pixels that had higher signal intensity during left eye monocular photic stimulation than during right eye simulation. The dimensions of the `activated' pixels are approximately 1 mm1 mm or slightly smaller and are reproducible. 98,99 The technique in this study was to rst use a binocular stimulation and then to use alternating left and right monocular stimulation. The data were, however, analyzed by looking for statistically signi cant differences in pixel intensities for only the monocular stimulation period, ignoring the binocular stimulation completely. The pixels identi ed as activated during either the left or right eye monocular stimulation also showed activation for the binocular period, as they should if they indeed represent ODCs as opposed to random statistical correlations. Figure 15b demonstrates ODCs in the human brain from a more recent study at 4 T by Menon and colleagues. 102 These images are obtained in the sagittal plane adjacent to the interhemispheric ssure and the ODCs are visualized as they appear on the cortical surface in this region of the visual cortex. The blue and red colors illustrate the columns associated with the two eyes. The columns are now seen along their long axis rather than in cross-section. These data demonstrate for the rst time that mapping of functional subunits in humans is possible in a noninvasive manner. The fmri time courses show that the hemodynamic response at the `hyperoxygenation phase' can be used at 4 T as a direct indicator of neuronal activity in cortical columns. This opens up the possibility of mapping specialized populations of neurons in humans that are not accessible to electrophysiological or other methods of invasive mapping. However, they do not yet resolve the issue of whether CBF increase coupled to increased neuronal activity and consequently BOLD response is spatially speci c and selective at the column level. The column images illustrated in Figure 15 were obtained using alternating monocular stimulation. Therefore, columns would be detectable as long as there was a difference in the deoxyhe- For list of General Abbreviations see end-papers

15 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 15 Figure 15 Detection of human ocular dominance columns (ODCs) during alternating monocular stimulation at 4 T. (a) Data obtained on a plane parallel to the calcarine ssure that intersects the ODCs perpendicular to their long axis. In this view, the ODCs should appear approximately as squares of 11 mm cross-section, separated by ~1 mm. There is a curving sulcus lined by the ODCs. (With permission from Menon et al. 99 ) (b) A more recent study at 4 T showing a sagittal plane adjacent to the interhemispheric ssure; the ODCs are visualized as they appear on the cortical surface. The blue and red colors illustrate the columns associated with different eyes. (With permission from Menon and Kim 101 and Goodyear and Menon 102 ) moglobin content of the inactive and active columns, and hence a difference in the BOLD effect, even if the deoxyhemoglobin content changed for both active and inactive columns, as illustrated in Malonek and Grinvald (Figure 14). 61 The results summarized above demonstrate that, despite the presence of several potential problems, fmri at 4 T has the speci city and sensitivity to map organizations in the millimeter or slightly smaller scale with the use of appropriate paradigms. This is unprecedented in human brain studies. We must emphasize the high- eld aspect of all of the afore-mentioned high-resolution studies; to date, the ODCs have not been detected at lower eld strengths. 5 SINGLE-TRIAL FUNCTIONAL MRI Most fmri studies utilize a `block' design where periods of a control state are interleaved with periods of task performance and/or sensory stimulation. The control period itself may also require the subject to perform a task and/or be subjected to sensory stimulation. The images are generated by examining the difference between the control and the tasking periods. Each of these periods are relatively long, typically a minute or more, and the subject executes the task many times. The picture that emerges from such a study is a time average that blurs important information regarding the temporal evolution of the neuronal activity in different parts of the brain. Equally important, cognitive effects such as learning, alteration in strategy, etc. that evolve with repeated executions are also averaged into the nal image generated. fmri is actually a realtime measurement. Single-slice fmri images can be acquired in tens of milliseconds, adequate to monitor neuronal responses. Unfortunately, the temporal response of fmri signals is dictated by the response of the vascular system, which is characterized with a time constant of seconds. This was demonstrated in one of the early fmri studies 103 and subsequently con rmed in numerous other reports. 67,104±108 However, even within this temporal regime, useful information on brain function can be obtained since numerous tasks and processes exist that necessarily engage the human brain for prolonged periods. In this temporal domain, signal changes in the fmri images track the temporal evolution of stimulation or mental task performance very well, albeit with a shift in time or a delay that lasts several seconds. This capability permits the acquisition of fmri data gated to a particular time point in stimulus onset or the instruction±execution sequence. In either case, this type of fmri data collection is referred to as event-related fmri. In event-related fmri, two distinct types of experiments have been performed. Buckner et al. acquired images gated to the onset of a task in a paradigm so that the temporal evolution of the fmri signal during and following the execution of the task could be temporally co-registered and averaged following repeated executions of the same task. 109 However, such aver- For References see p. 18

16 16 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY aging loses unique information associated with each execution of the task; subjects do not perform the same way each time because both brain function and performance is modulated by effects such as learning, alterations in strategy, errors, and habituation. In addition, when averages of single trials are performed, it is not possible to conclude much about the temporal differences that may be observed among different regions of the brain; such a difference was reported in the study by Buckner et al., where activation in left prefrontal areas involved in language were delayed ~2 s relative to extrastriate areas during a word generation task. 109 Of course, the human brain performs this task much faster than 2 s, and this time difference cannot really be attributed to differences in brain activation. Rather, differences must exist in the hemodynamic response of the fmri signal in different regions of the brain. This distinction, however, cannot be made from the data alone in this study, and the interpretation of the results would be ambiguous in less obvious cases when this approach is utilized. The second type of experiment was a true single-trial fmri study achieved with the use of high magnetic elds (4 T). 105,106,110±113 Intentionally and speci cally, it was demonstrated that (at least at 4 T) there exists enough sensitivity to monitor fmri signal evolution in a single execution of a task without averaging over many trials. This point is important. With this capability, it is then possible to perform many such single trial executions of a task and not average them but rather to store them separately and subsequently analyze and correlate the fmri data with differences in aspect of the subject's performance (e.g., response time, errors, etc.). In this way, the hemodynamic response time differences can also be factored out and distinguished from temporal behavior of neuronal activity. An alternative approach is to average such single trials but using performance or response criteria to pool together only those response that are similar. An example of a true single-trial study conducted at 4 T monitored the evolution of fmri signals in the human brain before, during, and subsequent to an instruction and execution of a motor task (Figure 16). Both the single-trial fmri signal intensity time courses in motor areas and the electromyograph (EMG) changes detected in the muscle are shown. The EMG showed no movements during the motor preparation period between Instruction and GO. Activation of the primary motor area and of the supplementary and premotor areas were observed during motor execution and preparation. These data are virtually identical to electrode recordings taken from the corresponding areas in a monkey cortex during execution of the same task, except that the fmri data are displaced in time by seconds relative to the actual time of neuronal activity. The data presented in Figure 16 also illustrate the sensitivity of the high- eld fmri method. Like the electrode recordings from primates, the fmri data demonstrate that V1 is active during the motor preparation period. This, however, has been a controversial issue with humans because position emission tomographic studies did not reveal activation during such motor preparation. 114 In the high- eld fmri study, this activation is detectable in a single execution of the task. This type of true single-trial experiment was also utilized to correlate aspects of activation with task performance. 106,113 The paradigm employed was the `mental rotation' task of Shepard and Metzler: 115 The subjects were presented with drawings of Figure 16 A single-subject, single-trial fmri (without averaging) study showing time courses in three regions of the brain together with the electromyograph (EMG) recording of muscle movement during a visually instructed delayed cued four- nger movement task. The presentation lasted 2.1 s and the GO signal was given at 9.1 s. M1, contralateral primary motor area; PM, bilateral premotor area; SMA, bilateral supplementary motor area. (With permission from Richter et al. 111 ) three-dimensional objects, examples of which are illustrated in Figure 17. In each task, a pair of objects were presented; they were either identical or mirror images and they were rotated relative to each other through varying degrees. The subject had to identify whether the pair was identical or a mirror image and report it by pressing one of two buttons. In this task, subject's response time depended on the angle through which the two objects were rotated relative to each other. The experiment started with the subject looking at similar but identical twodimensional objects. When ready, the subject commenced the scanning process by pressing a button. The three-dimensional objects were shown and the subject made a decision; after a suitable delay to allow for the hemodynamic response to return to basal levels, the process was repeated. Figure 18 displays signal intensity curves from the parietal lobe from two trials where the response time of the subject was different. The width of the response was evaluated for correct response only with respect to the response time. In each individual, linear correlation was found. However, the intercept corresponding to a response time of zero was signi cantly different for the different individuals, presumably re ecting fundamental differences in the hemodynamic response to elevated neuronal activity among individuals. When this subject-dependent variable was subtracted, each single-trial, single-subject data point for all subjects yielded an excellent correlation with response time (Figure 19). One potential confounding problem with such single-trial fmri studies is the presence of various types of spatiotemporal patterns in the fmri signals even under basal `resting' conditions. The T2 -weighted MRI signals uctuate with heart beat and respiration, although uctuations can be removed from the fmri data. 116±121 When cardiac and respiratory uctuations are suppressed, fmri signal from resting human brain still exhibits low-frequency oscillations at about 0.1 Hz. 121±124 Near-infrared For list of General Abbreviations see end-papers

17 FUNCTIONAL MRI AT HIGH FIELDS: PRACTICE AND UTILITY 17 Figure 17 The paradigm in a true single-trial `mental rotation' study. (With permission from Richter et al. 107 ) optical studies also show this slow oscillation. Therefore, techniques that can separate the fmri data into its various components (e.g., Mitra et al. 121 ) and in the process signi cantly improve effective SNR for detection of function will play a crucial role in mapping brain function, particularly in single-trial fmri studies with temporal resolution. 6 CONCLUSION Since its introduction, fmri has rapidly evolved to become the most signi cant method for investigating human brain function, and unique accomplishments have been realized at high magnetic elds. However, it must be realized that high magnetic eld MRI instruments are not optimized and re ned machines as clinical scanners. Presence of a high- eld magnet does not guarantee superior functional imaging results; instead, Figure 19 The width of the fmri response (in Figure 18) versus performance data from all subjects subsequent to removal of the time zero intercept. Each point represents a true single trial in a single subject (no averaging). (With permission from Richter et al. 107 ) advantages that are inherent in the high eld for functional mapping can easily be lost through instrumentation imperfections. Therefore, improvement in high- eld instrumentation are essential for future expansion of these fmri applications. Additional re nements in data collection schemes, motion correction, and statistical methods for data analysis, which are areas that are being actively pursued, will undoubtedly improve the already impressive capabilities of this methodology. 7 RELATED ARTICLES Figure 18 The T2 -weighted BOLD response in the parietal lobe for two different single trials (no averaging) for the paradigm outlined in Figure 17. The arrows indicate the time at which the subject responded. (With permission from Richter et al. 106 ) Hemodynamic Changes owing to Sensory Activation of the Brain Monitored by Echo-Planar Imaging; Image Processing of Functional MRI Data. For References see p. 18