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Insights into new techniques for high resolution functional MRI Seong-Gi Kim* and Seiji Ogawa† Non-invasive functional magnetic resonance imaging (fMRI) has opened a unique window into human and animal brain function, with a spatial resolution of a few millimeters and a temporal resolution of a few seconds. To further improve the current technical limitations of fMRI, various post-processing and data acquisition schemes were developed. Improved fMRI methods include variations of a conventional fMRI technique, mapping a single physiological parameter such as cerebral blood flow or cerebral blood volume, and direct mapping of neural activity. Advances in fMRI techniques allow scientists to map submillimeter columnar and laminar functional structures and to detect tens of millisecond neural activity in certain specific tasks. Addresses *Department of Neurobiology, University of Pittsburgh, E-1140 Biomedical Science Tower, 200 Lothrop Street, Pittsburgh, Pennsylvania 15261, USA; e-mail: [email protected] † Ogawa Laboratories for Brain Function Research, Hamano Life Science Research Foundation, 12 Daikyou-cho, Shinjuku-ku, Tokyo 160-0015, Japan; e-mail: [email protected] Current Opinion in Neurobiology 2002, 12:607–615 0959-4388/02/$ — see front matter © 2002 Elsevier Science Ltd. All rights reserved. Published online 14 August 2002 Abbreviations BOLD blood oxygenation level dependent CBF cerebral blood flow CBV cerebral blood volume fMRI functional magnetic resonance imaging S1 primary somatosensory cortex SNR signal-to-noise ratio

Introduction Following its introduction over a decade ago, functional magnetic resonance imaging (fMRI) using the blood oxygenation level dependent (BOLD) contrast has become the tool of choice for visualizing neural activity in the human brain [1]. The conventional BOLD approach has been extensively used for investigating various brain functions including vision, motor, language and cognition, with a spatial resolution ranging from a few millimeters to a few centimeters. Although this spatial scale is suitable to address the majority of cognitive and psychology issues, it is inadequate for studying functional parcellation at the millimeter or submillimeter level. To obtain high-resolution fMRI, two major issues have to be considered; first, the sensitivity; second, the specificity of the signal source (i.e. the spatial extent of the signal to the site of neuronal activity). Both factors determine the accuracy of the functional map. The importance of high sensitivity and specificity

For the effective utilization of fMRI techniques, it is essential to have high sensitivity. Often in fMRI studies,

higher resolution functional images appear more localized because of smaller statistically significant areas. This observation should be carefully examined as to whether it is coming from a lack of adequate signal-to-noise ratio (SNR) or from a genuine physiological phenomenon. Even if the source of the imaging signal is specific to neural activity, its technique may not be practically useful if its SNR is poor and signal averaging is not sufficient due to limited experimental time. High SNR of fMRI techniques is not sufficient for highresolution functional mapping if the signals that are being imaged do not have high specificity to the local neural activity. Thus, it is important to understand the signal source of BOLD fMRI and its fundamental limit of specificity, by examining a chain of events from neural activity to fMRI signals (reviewed in [2••,3,4,5•,6]). Increased neural activity induces an increase in metabolic demands. Thus, imaging metabolic changes (e.g. by 2-deoxyglucose autoradiography) will yield high spatial specificity, because such changes occur in the tissue at the site of the neuronal activation, and not in the vasculature. Because the early negative BOLD signal following the onset of the stimulus — referred to as a ‘dip’ — is likely to indicate an early oxygen consumption change [7], it attracted great attention. Changes in metabolism could modulate the hemodynamic responses, including cerebral blood flow (CBF), cerebral blood volume (CBV), and venous oxygenation levels. It has been well established that the magnitude of CBF change is well correlated with that of metabolic change and the extent of neural activation (e.g. [8] and see Figures 3 and 4 in [9]). Thus, CBF mapping can pinpoint the most active spot of neural activity even if the exact spatial extent of the CBF response is controversial [7,10••]. The most commonly used BOLD technique is sensitive to paramagnetic deoxyhemoglobin, which is located at the capillaries and the venous draining vascular system, reducing spatial specificity of the conventional gradient-echo BOLD signal (see below). Limits of specificity in hemodynamic fMRI

The intrinsic limit of spatial specificity of hemodynamicbased fMRI depends on how finely CBF is regulated. It has been suggested, on the basis of optical imaging studies, that CBF regulation is widespread beyond neuronally active areas [7]. However, recent studies suggest that intrinsic CBF changes are specific to submillimeter functional domains [10••]. This notion is supported by Harrison et al.’s observation that sphincter-like blood flow controlling systems are situated at or close to capillary branching points [11•]. The approximate distance between precapillary controlling valves and the start of venous drainage in the chinchilla is on the order of 100 to 150 µm [11•]; the average distance between capillary segments in

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Correlation between evoked potentials and the BOLD response. The results were obtained at 7 T in rat S1 during two stimuli with various inter-stimulus intervals. Right forepaw stimulation (indicated by the black up arrows) was followed by left forepaw stimulation. The observed evoked potentials (EPs) in the right S1 induced by the left paw stimulation (indicated by red down arrows) varied depending on the inter-stimulus intervals. Similarly, BOLD fMRI signals in the right S1 during two repeated stimuli (indicated by red bars) were dependent on duration of the inter-stimulus intervals. When the inter-stimulus interval

was 12.5 ms or 75 ms, but not 40 ms, significant evoked potentials were seen. The sharp spikes were generated by electric currents applied to the left paw during stimulation, indicating the actual time of the stimulus. fMRI signals closely followed the evoked potentials, indicating a tight coupling between neural activity and fMRI signal. When the inter-stimulus delay was 40 ms, both evoked potentials and BOLD fMRI signals decreased. This suggests that fMRI can detect temporal evolution of some neural activities among regions. Modified with permission from [14].

cats is 108 µm [12]. These values represent an ultimate limit of intrinsic spatial resolution for all hemodynamicbased neuroimaging techniques. The relatively high specificity of the hemodynamic signal is still, in principle, expected to be poor compared to the specificity of tissue metabolic signals. To obtain high spatial resolution in hemodynamic-based fMRI, capillary-based signals should be maximized while the large vessel contribution should be minimized. To further improve spatial and temporal resolution, non-hemodynamic, direct tissue signal changes should be sought.

in-plane resolution. In addition, we briefly discuss a new approach that aims to improve temporal resolution to neural time scales by using the BOLD technique.

In this review, we mainly discuss how to distinguish between ‘parenchyma’ and ‘large vessel’ components for each imaging method and how to improve spatial specificity and resolution of various imaging methods, including hemodynamic and non-hemodynamic techniques. Here, large vessels are defined as arteries, arterioles, venules and veins, which are larger than ~10 µm in diameter, and highresolution is defined as millimeter to submillimeter

Improvements in conventional BOLD approaches Tight correlation to neural activity and temporal constraints

To confirm a tight correlation between BOLD fMRI and neural activity at a supramillimeter spatial resolution, Logothetis et al. [13••] simultaneously measured conventional BOLD fMRI and neural activity in anesthetized monkeys. They found that the BOLD signal from a ~16 mm2 area is monotonically correlated with neural activity. Similarly, evoked potentials and hemodynamic responses are monotonically correlated [2••,5•,14,15•]. These studies indicate that BOLD fMRI is well correlated with neural activity at a coarse (several millimeters) spatial resolution. Using this property, Ogawa et al. [14] devised a novel approach to detect neural interaction on the millisecond time scale using fMRI. Figure 1 shows an example of

Insights into new techniques for high resolution functional MRI Kim and Ogawa

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Figure 2 Functional images of conventional BOLD, CBF and synaptic activities from one anesthetized rat obtained at 9.4 T. The maps of forepaw somatosensory stimulation are overlaid on respective basal images. (a) An anatomical coronal image of the rat brain, showing the region of interest expanded in each of the three fMRI maps, is represented (top left). To compare between the three activation maps — (b) BOLD, (c) manganese-enhanced and (d) CBF — 50 pixels with the highest statistical values were color-coded. A cartoon showing a section of rat S1 (layers I to VI; delineated by blue lines) is also displayed in (b)–(d). The colored bar indicates the values of percentage changes, ranging from 1 to 5% (red to yellow) for the conventional BOLD (b) map and 50 to 150% (red to yellow) for the CBF (d) map. For the synaptic activity map (c), manganese ions (Mn2+) were used as a calcium analogue as well as an MRI contrast agent. In (c) and (d) activation maps, the ‘hot’ spot indicated by yellow was located in layer IV of S1. This

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detecting neural interaction between the rat primary somatosensory cortex (S1) of two hemispheres, using fMRI with submillimeter spatial resolution. Right paw somatosensory stimulation was followed by left paw stimulation with various inter-stimulus intervals. When the inter-stimulus interval was zero or 12.5 ms, both stimulation-evoked potentials and fMRI signals induced by the left paw stimulation were observed in the right S1. But, with an inter-stimulus interval of 40 ms, both signals disappeared. This indicates that the right S1 activity induced by the left paw stimulation is canceled by inhibitory neural activity from the preceding right paw stimulation [14]. Thus, fMRI can be used to detect neural temporal interaction between regions. This will be a very valuable method to examine the temporal evolution of certain neural events among regions by controlling input tasks. It should be noted that the temporal resolution of conventional fMRI approaches is on the order of a few hundred milliseconds to a few seconds (see [16] for a recent review in this journal). Spatial constraints

It is well known that, with typical fMRI acquisition parameters, the BOLD response is particularly sensitive in and around large draining veins, which are usually distant from the sites of elevated neuronal activity. It is therefore a reasonable assumption that conventional BOLD-based high-resolution fMRI may misrepresent the functionally less specific large vessel contribution. To examine the spatial extent of conventional BOLD fMRI, Disbrow et al. [17] measured BOLD fMRI and electrophysiological activity in anesthetized monkeys at 1.5 T. They found that the largest mismatch between the two measurements was located at areas close to large vessels, as expected. Thus, it is desirable to remove or minimize large vessel contributions

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using post-processing techniques and improved data acquisition methods. Post-processing approaches detect large draining vessels by using various BOLD signal characteristics such as a large intensity [18••], a delayed response, and a significant phase change [19]. Although these subjective criteria are effective, it may not be sufficient to detect and remove all large vessel contaminations. Remaining ‘common’ large vessel contributions to the BOLD signal can be removed by the differential imaging approach, if functional territories are complementary during orthogonal stimuli [18••,20•]. However, this subtraction method cannot be used in most fMRI studies because of unknown orthogonal stimulation conditions. Thus, it is critical to remove or minimize the draining vessel contribution. Several improved data acquisition methods can be used to minimize large vessel contributions [18••,21•]. The early negative BOLD signal can be utilized at high magnetic fields because the magnitude of the dip increases at higher magnetic fields. As the source of the early negative BOLD signal is likely induced by an increase in oxygen consumption without a similar CBF response [7], the early negative BOLD response is indicative of metabolic changes and provides high spatial specificity [22–25]. However, the dip signal does not originate directly from tissue oxygenation changes, but from deoxyhemoglobin concentration changes in the blood, induced by increased oxygen consumption. Similarly to conventional positive BOLD signals, the dip initially occurs in the parenchyma and drains into the venous blood system including large veins. Thus, only the early portion of the negative BOLD signal is specific to the site of neuronal activity increases. To detect the early response and utilize the high spatial

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CBF-weighted functional images. The images were obtained during (a) left finger movements in humans and during (b) single orientation visual stimulation in anesthetized cats. (a) To obtain the perfusion contrast, flow direction (indicated by an up arrow) has to be considered. Thus, 5 mm-thick transverse planes were selected for CBF-based fMRI studies. The background image was perfusionweighted; higher signal areas have higher inflow rates. Functional activity areas are located in the gray matter of the contralateral primary motor cortex (indicated by a red arrow). Interestingly, no large signal changes were observed at the edge of the brain, an artefact that is often seen in conventional BOLD functional maps. A, anterior; R, right hemisphere. (b) A statistical t-value map of one animal and the functional map thresholded at p < 0.05 of another animal are shown here. Statistical t-values were calculated by comparing intensities of images obtained during stationary and moving gratings on a pixel-bypixel basis. Images were obtained in a tangential plane of the visual cortex, area 18. The top of the images is anterior and arrows indicate the midline. Both maps were obtained without using any subtraction method. Clearly, localized CBF responses can seen at submillimeter spatial resolutions (orange areas in right panel), demonstrating the improvement of fMRI signals. Further, irregular shaped active patchy clusters are separated by ∼1 mm and are ∼0.5 mm wide, indicating that the CBF response is specific to active orientation columns [10••]. The gray scale bar represents t values from 0 to 4.

specificity of the dip signal, MR images need to be acquired with high spatial and temporal resolution, which induces significant reductions of SNR [26]. It should be noted that the duration and magnitude of the dip in anesthetized cats are closely related to physiological conditions including anesthesia level, end-tidal CO2, and blood pressure [27]. Consequently, the early dip is highly susceptible to physiological fluctuations as well as basal conditions and has a small overall SNR [10••]. Thus, the dip is not likely to be viable for high-resolution fMRI studies in humans.

At higher magnetic fields, the large vessel contribution decreases while the tissue contribution increases due to the quadratic dependence of tissue signals on magnetic field strength [28]. With higher SNR and better specificity, high magnetic fields (≥3 T) are an effective approach to reduce large vessel contributions to the BOLD signal and obtain high-resolution images. By combining post-processing approaches, human ocular dominance columns were convincingly mapped at 4 T [18••,20•]. However, even at an ultrahigh field of 9.4 T, the conventional BOLD technique contains the extravascular contribution around large vessels (the yellow pixels at the edge of the brain in Figure 2b) [29]. To further improve spatial specificity, spin-echo BOLD fMRI at high fields can be used [28]. The difference between conventional gradient-echo and spin-echo BOLD techniques is that the spin-echo technique uses one additional radiofrequency pulse (which is the same as T2-weighting in diagnostic imaging). This additional pulse minimizes the extravascular BOLD effect around large draining vessels. Subsequently, parenchyma signals and the intravascular component in large vessels remain. The intravascular signal can be removed by using high magnetic fields such as 7 T or other flow-removing methods used in diffusion imaging [29]. However, the drawbacks of the spin-echo BOLD technique are inferior sensitivity and temporal resolution, when compared to those of the commonly used gradient-echo BOLD. Thus, high-resolution tissue-specific spin-echo BOLD images cannot be easily obtained at 1.5 T because of low sensitivity. Disadvantages of higher fields include increases in power deposition induced by radiofrequency pulses, the dielectric effect, and susceptibility around sinus and air-tissue boundaries. If the CBF response is widespread, hot spots of parenchyma BOLD signals may not colocalize with those of CBF and metabolic images at high spatial resolution [18••], because the BOLD signal is related to a mismatch between CBF and oxygen consumption changes. To investigate this issue, we obtained spin-echo BOLD fMRI data of the rat somatosensory cortex at 9.4 T during somatosensory stimulation, and found that laminar-specific functional imaging signals can be detected [29]. The correlation between spin-echo BOLD signals and the CBF response is excellent on a pixel-by-pixel basis, suggesting that spinecho BOLD signals at very high fields are a good index of the CBF response [30•]. This observation suggests that a higher tissue-level BOLD signal (without draining vein contribution) indicates higher neural activity.

Cerebral blood flow techniques Alternative techniques to the BOLD approach, which are sensitive mainly to parenchyma, can be considered for obtaining complementary information regarding brain activity. For this, CBF measurements using 15O-labeled water as an exogenous positron emission flow tracer have been extensively performed for functional mapping [8,9].

Insights into new techniques for high resolution functional MRI Kim and Ogawa

Similarly, CBF-weighted functional images can be obtained by employing arterial blood water as an endogenous flow tracer with high spatial and temporal resolution. Labeled water is prepared by perturbing the magnetic status of arterial water using a radiofrequency pulse [31]. Then, magnetically labeled water moves into the capillaries in the imaging slice and exchanges with tissue water. Because a background water signal exists, it is important to remove non-flow related water signals. Thus, two images are typically acquired; one with arterial magnetic labeling and the other without labeling [31–33]. The difference between the two images is directly related to CBF, and relative CBF changes due to physiological perturbations can be measured. Obviously, during an arterial labeling time, labeled water will fill up the arterial vasculature before reaching the capillaries and the tissue. Thus, it is essential to have sufficient labeling time; for example, >0.5 s in humans and >0.2 s in rats [34]. Following this delivery time of labeled spins to the capillaries, the CBF contrast mostly reflects perfusing water spins that have permeated the capillary walls and entered the extravascular space. Large draining vessels do not significantly contribute to CBF-weighted MRI signals due to the relatively short half-life of labeled water (1–2 s depending on magnetic fields). Figure 3 shows CBF-weighted fMRIs in humans and animals. A background image is a CBF-weighted difference image described above (Figure 3a); gray matter areas have higher signal intensity than white matter due to higher CBF. Further, because the relative CBF change is linearly correlated to a metabolic change, the CBF change is a good index of the magnitude of metabolic/neural response. The highest CBF change was observed in layer IV of the rat S1 during somatosensory stimulation (see Figure 2d) [30•] and active columns in cat area 18 during visual stimulation (see Figure 3b) [10••]. The specificity of CBF-based fMRI signals is better than that of conventional BOLD contrast, but comparable to that of the high field spin-echo BOLD technique [30•]. Proper CBF contrast is only achieved when enough time is allowed for the labeled arterial water to travel into the region of interest and exchange with tissue water. This makes it difficult to detect changes in CBF with a temporal resolution greater than the lifetime of the labeled water. Acquisition of a pair of images can further reduce temporal resolution and consequently SNR. Also, the selection of CBF-based imaging slices is dictated by the direction of flow. Thus, CBF-based imaging techniques have not been widely used for routine fMRI studies. It should be noted that when not considering temporal resolution, the SNR of CBF-based fMRI can be as good as that of conventional BOLD [35]. Acquisition of a pair of flow-weighted and control images is necessary to obtain quantitative CBF changes. However, to achieve CBF-weighted qualitative fMRI contrast, flow-weighted images alone can be

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obtained, resulting in improvement of SNR and temporal resolution [36,37•].

Contrast-enhanced cerebral blood volume techniques The CBV-based technique uses an exogenous intravascular contrast agent with high susceptibility. Following a bolus injection of a paramagnetic agent [38], relative CBV can be determined by integrating MRI signals over the first pass of the contrast agent. Although measurement of CBV by this technique has been used to perform the first functional images in humans during visual stimulation [38], its utility is severely limited by the need for repeated bolus injections separated by long intervals. To overcome this problem, iron oxide particles have been used as a long half-life intravascular contrast agents [39–42], similar to endogenous irons in deoxyhemoglobin molecules (used for the BOLD contrast). The typical half-life of the iron oxides in blood is approximately 4 h in rats [40] and >8 h in non-human primates [43••]. During steady-state conditions, established following the infusion of an intravascular contrast agent, an increase in CBV during stimulation will induce an increase in intravoxel contrast agents, and consequently a decrease in the MRI signal. Although the change in deoxyhemoglobin (used in the BOLD signal) will contribute to the contrast-agentenhanced MRI signal, exogenous contrast agents with sufficiently high doses dominate the environment surrounding deoxyhemoglobin. Thus, fMRI following the injection of contrast agents is predominantly weighted by the CBV change. In typical fMRI studies, neural activity induces BOLD signal increases, whereas CBV-weighted MRI results in signal decreases. This CBV approach has two major advantages over the conventional BOLD technique: spatial specificity and SNR. Because changes in CBV are well correlated with CBF changes [44–48] and the CBF response appears to be specific to tissue, it is expected that CBV is also specific to parenchyma. Small vessels including precapillary arterioles dilate rigorously during neural stimulation, whereas large vessels do not dilate much [47]. Because the MRI signal is very weak in a voxel containing large vessels as well as in its immediate vicinity (due to a large amount of contrast agent), any change induced by a decrease in deoxyhemoglobin may not be detected. This is the same mechanism as used by high magnetic field BOLD for reducing the large vascular contribution. Thus, the spatial specificity of CBV at low magnetic fields is similar to that of the conventional BOLD contrast at ultrahigh fields such as 9.4 T. Figure 4 shows a comparison between conventional BOLD and CBV-weighted fMRI in the cat visual cortex at 4.7 T. Clearly, contrast-agent-enhanced CBV-weighted signals are located in the middle of the cortex, whereas the largest signal change in conventional BOLD occurs at the surface of the cortex and between the two hemispheres. This demonstrates the improvement in spatial specificity when using contrast agents.

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The SNR of the CBV measurement is dependent on the dose of the contrast agent, the imaging parameters, and the magnetic field. Because the BOLD contribution is supralinearly (between one and two) dependent on magnetic field [28], whereas the CBV change remains constant and iron oxides are saturated at high fields [40], the SNR gain of the CBV technique over the BOLD technique is higher at low fields. At 1.5–2 T, the improvement in sensitivity is ~five-fold [40,43••], while the gain is ~1.5–2.0-fold at 4.7 T [40]. At ultrahigh fields (e.g. 9.4 T), the SNR gain is expected to be minimal. One obvious concern is whether this contrast agent is safe for long-term use. Behaving monkeys during repeated injections of contrast agents, studied for up to 10 months, did not show any obvious negative health effects [43••], suggesting that contrast agents can be used repeatedly. To detect CBV changes and increase sensitivity, a dose of contrast agent should be sufficiently high — typically 5–10 mg iron per kg body weight for animal studies. In human studies with a much smaller dose (~1 mg/kg), the CBV change was not detected during neural stimulation [49]. Thus, the major problem of the long half-life contrast-agent-based fMRI approach for human studies is that it requires a large amount of exogenous contrast agent. Taken together, the CBV-based technique is a highly sensitive hemodynamic-based fMRI technique for animal studies.

Non-hemodynamic techniques The above-mentioned fMRI techniques are based on hemodynamic responses induced by neural activity. To further improve spatial resolution and specificity, it would be ideal to detect signal changes in tissue induced by neural activity directly. Here, we review three direct measurement techniques: water diffusion measurement, direct imaging of electric current effect, and calcium influx measurements using manganese as a tracer. Recently,

Conventional BOLD and contrast-enhanced CBV-weighted fMRI of cat visual cortex during visual stimulation, overlaid on an anatomic image. Both images were obtained in a coronal plane with 150 × 150 µm2 in-plane spatial resolution and 2 mm slice thickness at 4.7 T. CBV-weighted images were obtained after the intravascular injection of iron oxide contrast agents (7 mg iron/kg body weight). During visual stimulation, the BOLD signal increases (red:yellow ratio) while the CBV-weighted signal decreases (blue:purple ratio), indicating an increase in CBV. The largest BOLD signal change was observed on the surface of the cortex, whereas the largest CBV change was located in the middle of the cortex. This demonstrates the improvement in spatial specificity of the contrast-enhanced CBV technique. Color bars represent t-values from 1 to 10. D, dorsal; L, lateral. Images courtesy of Noam Harel at the University of Minnesota.

Darquie et al. [50•] reported that a decrease in tissue water diffusion was observed during visual stimulation; this was possibly due to swelling of the neurons and glial cells. A similar decrease in diffusion was observed during seizures and cortical electric shocks [51]. Although the water diffusion technique can improve spatial specificity and is extremely promising, the time course of the water diffusion change is similar to that of typical hemodynamic responses. Thus, there is concern that the diffusion change may relate to the susceptibility change induced by deoxyhemoglobin [52]. Further studies are necessary to examine the source of the diffusion change. It should be noted that the SNR of this measurement is inferior to conventional BOLD contrast. The electric current associated with neural activation can produce a local magnetic field that is the source of extraskull magnetometry or MEG. The magnetic resonance detection of the current-induced magnetic field has been a topic for some laboratories [53–55]. On the basis of in vitro phantom studies with electrical current conductors, a 10 µA (~2 nT) current can be detected by observing phase changes in MRI signals [54,55]; a 90 nA current was detected by using a subtraction of two images obtained with opposite gradient polarity, but with the same susceptibility effect [53]. Using the subtraction approach, Kamei et al. [53] reported that direct neuronal mapping has been achieved in humans without any contamination of the BOLD signal. However, this approach is still too premature for practical applications, due to its poor sensitivity. Further studies are necessary for confirmation of this important initial report and for improvement in SNR. Koresky’s laboratory used manganese ions as an in vivo, transsynaptic, anterograde, neuronal tract-tracer as well as a paramagnetic contrast agent [56]. This approach has

Insights into new techniques for high resolution functional MRI Kim and Ogawa

been successfully demonstrated to trace an optical tract from the retina to the thalamus and superior colliculus, and an olfactory tract from the naris to the olfactory bulb [56–58]. In a manner similar to this approach, manganese ions, used as a calcium analogue as well as an MRI contrast agent, can be used to visualize calcium influx from extracellular into intracellular compartments through voltage-gated calcium channels [59,60]. The preferential accumulation of paramagnetic manganese ions in actively firing neurons, associated with increased neuronal activity, results in the marked change of the MRI signal (see Figure 2c). Because the manganese ions that accumulate in the intracellular space have a half-life of two to three months, they can be visualized at later time points using MRI. This allows for tasks/stimulations to be performed outside the MRI environment. Although calcium influx is instantaneous compared to hemodynamic responses, this fast process cannot be imaged because of insufficient accumulation of manganese ions during a short period of time. Also, manganese ions cannot penetrate into the extravascular space without breaking the blood–brain barrier, which requires the use of drugs such as mannitol. In addition to concerns regarding toxicity, manganese ions accumulate non-specifically at the choroid plexus, an area with a high permeability to endothelial cells.

Conclusions Many techniques for high spatial and temporal resolution fMRI have been proposed and developed over the last decade. For the ever-growing neuroscience application of fMRI, both SNR and accuracy of mapping signals should be considered. On many occasions, MRI techniques with superb specificity do not provide adequate SNR for routine experiments. Especially, in human studies, the improvement of SNR by signal averaging is limited due to a relatively short experimental time. Because of its high sensitivity, the most widely used conventional BOLD fMRI technique is the method of choice for supramillimeter functional mapping and for high temporal resolution studies in humans. To further increase the sensitivity and spatial specificity of the BOLD signals, the use of a high magnetic field can be an expensive, but relatively simple solution. This is the main reason that many institutes are rushing to purchase high-field MRI systems. Caveats to the use of ultrahigh fields, such as 7 T, are the technical demands; ultrahigh fields are not a solution for all fMRI studies. CBF and CBV-weighted fMRI techniques are very promising approaches because they represent single physiological parameters and are specific to the parenchyma; however, the BOLD signal, which is the convolution of several components, may not be specific to tissue. CBF and CBV approaches provide higher spatial resolution than the conventional BOLD method. Although it cannot be easily applied to humans, the contrast-enhanced CBVweighted fMRI technique also provides high SNR. Using improved hemodynamic-based fMRI technologies, submillimeter functional structures can be mapped in animals

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as well as in humans. Non-hemodynamic fMRI mapping methods are still in their infancy with regard to human applications and require further improvements. In the future, chemists may be able to synthesize safe and powerful MRI tracers, which would be directly sensitive to neural activity; MR physicists may devise new approaches that are more specific and highly sensitive to neuronal activity. Current and newly improved MRI technologies allow neuroscientists to map laminar-specific and columnarspecific functional architectures.

Acknowledgements Supported by the National Institutes of Health (NS38295, NS40719, MH57180), the McKnight Foundation and the Keck Foundation. We thank all members of the Center for Magnetic Resonance Research, especially Noam Harel, Hiro Fukuda and Essa Yacoub at the University of Minnesota for exciting discussions.

References and recommended reading Papers of particular interest, published within the annual period of review, have been highlighted as:

• of special interest •• of outstanding interest 1.

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Lauritzen M: Relationship of spikes, synaptic activity, and local changes of cerebral blood flow. J Cereb Blood Flow Metab 2001, 21:1367-1383. This is an excellent review article. Lauritzen reports on the relationship between neural activity and hemodynamic responses, using the rat cerebellar cortex as a model. Spike activity does not affect the hemodynamic response (see his Figure 4), whereas synaptic activity predominantly induces cerebral blood flow changes. The magnitudes of CBF change and synaptic field potentials have a monotonic relationship (see Figures 5, 6, 8) and a linear relationship in some cases. This observation is consistent with studies in the monkey visual cortex [13••]. 3.

Attwell D, Laughlin SB: An energy budget for signaling in the grey matter of the brain. J Cereb Blood Flow Metab 2001, 21:1133-1145.

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Bandettini PA, Ungerleider LG: From neurons to BOLD: new connections. Nat Neurosci 2001, 4:864-866.

5. Arthurs OJ, Boniface S: How well do we understand the neural • origin of the fMRI BOLD signal? Trends Neurosci 2002, 25:27-31. This is a good review article, describing our current understanding of and the controversies surrounding the neurophysiological source of BOLD fMRI. 6.

Heeger DJ, Ress D: What does fMRI tell us about neuronal activity? Nat Rev Neurosci 2002, 3:142-151.

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Malonek D, Grinvald A: Interactions between electrical activity and cortical microcirculation revealed by imaging spectroscopy: implication for functional brain mapping. Science 1996, 272:551-554.

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Raichle ME: Circulatory and metabolic correlates of brain function in normal humans. In Handbook of Physiology — The Nervous System, vol 5. Edited by the American Physiological Society. Bethesda: American Physiological Society; 1987:643-674.

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Seitz RJ, Roland PE: Vibratory stimulation increases and decreases the regional cerebral blood flow and oxidative metabolism: a positron emission tomography (PET) study. Acta Neurol Scand 1992, 86:60-67.

10. Duong TQ, Kim D-S, Ugurbil K, Kim S-G: Localized cerebral blood •• flow response at submillimeter columnar resolution. Proc Natl Acad Sci USA 2001, 98:10904-10909. Here, the spatial specificity of CBF regulation is investigated using a cat orientation column as a model. The authors used arterial spin labeling MRI technique in cat area 18 during single orientation stimulation. They find that functional maps obtained using this new approach are immune to large draining vessels and specific to parenchyma. Further, functionally active areas appear like columnar structures based on the average distance of patchy clusters. The most important finding is that the hemodynamic

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response is regulated at a submillimeter columnar level, suggesting that columnar structures can be obtained without using differential imaging methods if large vessel contributions are minimized (see also [11•,18••,20•]). 11. Harrison RV, Harel N, Panesar J, Mount R: Blood capillary • distribution correlates with hemodynamic-based functional imaging in cerebral cortex. Cereb Cortex 2002, 12:225-233. Harrison et al. compare the spatial distribution of the auditory stimulus response from intrinsic optical images in the auditory cortex to the distribution of blood vessels, visualized in the same animal by corrosion casts. They find that intrinsic signals are well correlated with capillary densities. Further, blood-controlling sphincters are located at capillary branching points in the arterioles (see their Figure 7), suggesting that each capillary controls flow independently. Thus, the ultimate spatial resolution of the hemodynamic response is the spatial extent of a single capillary (see their Figure 8). This is consistent with the observation of Duong et al. [10••]. 12. Pawlik G, Rackl A, Bing RJ: Quantitative capillary topography and blood flow in the cerebral cortex of cats: an in vivo microscopic study. Brain Res 1981, 208:35-58. 13. Logothetis NK, Pauls J, Augath M, Trinath T, Oeltermann A: •• Neurophysiological investigation of the basis of the fMRI signal. Nature 2001, 412:150-157. This is the first demonstration of simultaneous BOLD and multiunit recordings in the same animal. This paper provides rich information regarding the neurophysiological source of the BOLD signal and shows a great technical achievement. The authors demonstrate that the BOLD fMRI signal is closely related to local field potentials rather than to spike activity, similar to findings from Lauritzen’s laboratory and others [2••,14,15•]. However, close examination of Figure 5 in this paper shows that local field potentials and spike activity behave similarly relative to the BOLD signal change, suggesting that the BOLD signal can also be an index of spike activity. Another interesting observation is that the BOLD signal change appears significant even when normalized local field potentials and spike activity approach zero. This observation indicates that a sigmoidal relationship between neural activity and the hemodynamic response is expected (see also [15•]). 14. Ogawa S, Lee T-M, Stepnoski R, Chen W, Zhu X-H, Ugurbil K: An approach to probe some neural systems interaction by functional MRI at neural time scale down to milliseconds. Proc Natl Acad Sci USA 2000, 97:11026-11031.

21. Yacoub E, Shmuel A, Pfeuffer J, Van De Moortele P, Adriany G, • Andersen P, Vaughan J, Merkle H, Ugurbil K, Hu X: Imaging brain function in humans at 7 Tesla. Magn Reson Med 2001, 45:588-594. Here, human fMRI at 7T is reported for the first time. The sensitivity and spatial specificity of the conventional BOLD signal was improved using higher magnetic fields. To take advantage of the high SNR achieved by high fields, signal fluctuations by physiological motions should be corrected because fluctuations also increase proportionally. 22. Kim D-S, Duong TQ, Kim S-G: High-resolution mapping of iso-orientation columns by fMRI. Nat Neurosci 2000, 3:164-169. 23. Cannestra AF, Pouratian N, Brookheimer SY, Martin NA, Becker DP, Toga AW: Temporal spatial differences observed by functional MRI and human intraoperative optical imaging. Cereb Cortex 2001, 11:773-782. 24. Mayhew J, Johnston D, Martindale J, Berwick J, Zheng Y: Increased oxygen consumption following activation of brain: theoretical footnotes using spectroscopic data from barrel cortex. Neuroimaging 2001, 13:975-987. 25. Yacoub E, Shmuel A, Pfeuffer J, Van De Moortele P, Adriany G, Ugurbil K, Hu X: Investigation of the initial dip in fMRI at 7 Tesla. NMR Biomed 2001, 14:408-412. 26. Buxton RB: The elusive initial dip. Neuroimaging 2001, 13:953-958. 27.

Harel N, Lee S-P, Nagaoka T, Kim D-S, Kim S-G: Origin of negative blood oxygenation level dependent fMRI signals. J Cereb Blood Flow Metab 2002, 22:708-717.

28. Ogawa S, Menon RS, Kim S-G, Ugurbil K: On the characteristics of functional magnetic resonance imaging of the brain. Annu Rev Biophys Biomol Struct 1998, 27:447-474. 29. Lee S-P, Silva AC, Ugurbil K, Kim S-G: Diffusion-weighted spinecho fMRI at 9.4 T: microvascular/tissue contribution to BOLD signal change. Magn Reson Med 1999, 42:919-928.

15. Nielsen AN, Lauritzen M: Coupling and uncoupling of activity • dependent increase of neuronal activity and blood flow in rat somatosensory cortex. J Physiol 2001, 533:773-785. These authors investigate the relationship between neural activity and CBF response in S1 during somatosensory stimulation using laser Doppler flowmetry and field potential measurements. Consistent with their previous cerebellar studies (see [2••]), a monotonic relationship between field potentials and CBF response was observed. Also, layer-specific CBF changes were seen, suggesting that layer-specific functional mapping is possible using hemodynamic-based fMRI techniques.

30. Lee S-P, Silva AC, Kim S-G: Comparison of diffusion-weighted • high-resolution CBF and spin-echo BOLD fMRI at 9.4 T. Magn Reson Med 2002, 47:736-741. This study shows the quantitative relationship between CBF and tissuespecific spin-echo BOLD changes in rat S1. When the large vessel contribution to the BOLD signal is minimized using the spin-echo technique at 9.4 T, the magnitude of tissue-specific BOLD signals is indicative of the magnitude of CBF responses on a pixel-by-pixel basis (see also [29]). This finding augments the theory that the CBF response is not widespread, but localized to submillimeter functional structures. On the basis of this study, human ocular dominance columns can be correctly assigned from differential maps (see discussion in [18••]); a higher BOLD signal area indicates a higher CBF/neural activity area.

16. Menon RS: Imaging function in the working brain with fMRI. Curr Opin Neurobiol 2001, 11:630-636.

31. Detre JA, Leigh JS, Williams DS, Koretsky AP: Perfusion imaging. Magn Reson Med 1992, 23:37-45.

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32. Edelman RR, Siewert B, Darby DG, Thangaraj V, Nobre AC, Mesulam MM, Warach S: Qualitative mapping of cerebral blood flow and functional localization with echo-planar MR imaging and signal targeting with alternating radio frequency. Radiology 1994, 192:513-520.

Disbrow EA, Slutsky DA, Roberts TPL, Krubitzer LA: Functional MRI at 1.5 Tesla: a comparison of the blood oxygenation leveldependent signal and electrophysiology. Proc Natl Acad Sci USA 2000, 97:9718-9723.

18. Cheng K, Waggoner R, Tanaka K: Human ocular dominance •• columns as revealed by high-field functional magnetic resonance imaging. Neuron 2001, 32:359-374. Here, human ocular dominance columns were mapped using the conventional BOLD technique with post-processing approaches at 4 T. Although not the first paper to map human ocular dominance columns using MRI, the systematic analyses and high reproducibility described in Figure 9 of this paper demonstrate that ocular dominance columns obtained using the differential imaging method are genuine. The discussion section provides an excellent reference for high-resolution fMRI studies. This study, along with studies from Menon’s laboratory [20•], shows that high magnetic fields can be used for mapping high-resolution functional images at columnar resolution. However, this technique still relies on the subtraction method (see also [10••,20•]). 19. Menon RS: Postacquisition suppression of large-vessel BOLD signals in high-resolution fMRI. Mag Reson Med 2002, 47:1-9. 20. Goodyear B, Menon RS: Brief visual stimulation allows mapping of • ocular dominance in visual cortex using fMRI. Hum Brain Mapp 2001, 14:210-217. Goodyear and Menon obtain human ocular dominance columns using conventional BOLD fMRI and post-processing methods at 4 T, similarly to [18••]. They demonstrate that early BOLD responses are more specific than late BOLD responses. See also [10••].

33. Kim S-G: Quantification of relative cerebral blood flow change by flow-sensitive alternating inversion recovery (FAIR) technique: application to functional mapping. Magn Reson Med 1995, 34:293-301. 34. Calamante F, Thomas D, Pell G, Wiersma J, Turner R: Measuring cerebral blood flow using magnetic resonance imaging techniques. J Cereb Blood Flow Metab 1999, 19:701-735. 35. Aquirre G, Detre JA, Zarahn E, Alsop D: Experimental design and the relative sensitivity of BOLD and perfusion fMRI. Neuroimaging 2002, 15:488-500. 36. Kwong KK, Belliveau JW, Chesler DA, Goldberg IE, Weisskoff RM, Poncelet BP, Kennedy DN, Hoppel BE, Cohen MS, Turner R et al.: Dynamic magnetic resonance imaging of human brain activity during primary sensory stimulation. Proc Natl Acad Sci USA 1992, 89:5675-5679. 37. Duyn J, Tan CX, van Gelderen P, Yongbi M: High-sensitivity single • shot perfusion-weighted fMRI. Mag Reson Med 2001, 46:88-94. In this paper, Duyn et al. describe an eloquent technique to improve the sensitivity of the CBF-weighted fMRI technique. By suppressing background water signals using radiofrequency pulses, temporal resolution and sensitivity can be gained over the conventional CBF-weighted fMRI techniques.

Insights into new techniques for high resolution functional MRI Kim and Ogawa

38. Belliveau JW, Kennedy DN, McKinstry RC, Buchbinder BR, Weisskoff RM, Cohen MS, Vevea JM, Brady TJ, Rosen BR: Functional mapping of the human visual cortex by magnetic resonance imaging. Science 1991, 254:716-719. 39. Kennan RP, Scanley BE, Innis RB, Gore JC: Physiological basis for BOLD MR signal changes due to neuronal stimulation: separation of blood volume and magnetic susceptibility effects. Magn Reson Med 1998, 40:840-846. 40. Mandeville JB, Marota JJ, Kosofsky BE, Keltner JR, Weissleder R, Rosen BR: Dynamic functional imaging of relative cerebral blood volume during rat forepaw stimulation. Magn Reson Med 1998, 39:615-624. 41. van Bruggen N, Busch E, Palmer JT, Williams S-P, de Crespigny AJ: High-resolution functional magnetic resonance imaging of the rat brain: mapping changes in cerebral blood volume using iron oxide contrast media. J Cereb Blood Flow Metab 1998, 18:1178-1183. 42. Mandeville JB, Jenkins BG, Kosofsky BE, Moskowitz MA, Rosen BR, Marota JJA: Regional sensitivity and coupling of BOLD and CBV changes during stimulation of rat brain. Magn Reson Med 2001, 45:443-447. 43. Vanduffel W, Fize D, Mandeville JB, Nelissen K, van Hecke P, •• Rosen BR, Tootell RBH, Orban GA: Visual motion processing investigated using contrast agent-enhanced fMRI in awake behaving monkeys. Neuron 2001, 32:565-577. This is the first study using the CBV-weighted fMRI technique at 1.5 T in behaving monkeys. It demonstrates that intravascular contrast agents — iron oxides — are safe for long-term usage. The improvement of sensitivity using contrast agents was clear. This opens a new window for animal fMRI studies at medium magnetic fields (1.5T) with high sensitivity and specificity, similar to doing BOLD experiments at extremely high fields such as 9.4 T (see also [39–41]). This straightforward approach does not require expensive high field magnets. The MRI technique in itself is the same as the BOLD technique, except for the injection of contrast agents; thus, this CBV-weighted fMRI does not need sophisticated MRI techniques. 44. Grubb RL, Raichle ME, Eichling JO, Ter-Pogossian MM: The effects of changes in PaCO2 on cerebral blood volume, blood flow, and vascular mean transit time. Stroke 1974, 5:630-639. 45. Jones M, Berwick J, Johnston D, Mayhew J: Concurrent optical imaging spectroscopy and laser-Doppler flowmetry: the relationship between blood flow, oxygenation, and volume in rodent barrel cortex. Neuroimaging 2001, 13:1002-1015. 46. Ito H, Takahashi K, Hatazawa J, Kim S-G, Kanno I: Changes in human regional cerebral blood flow and cerebral blood volume during visual stimulation measured by positron emission tomography. J Cereb Blood Flow Metab 2001, 21:608-612. 47.

Lee S-P, Duong T, Yang G, Iadecola C, Kim S-G: Relative changes of cerebral arterial and venous blood volumes during increased cerebral blood flow: implications for BOLD fMRI. Magn Reson Med 2001, 45:791-800.

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48. Jones M, Berwick J, Mayhew J: Changes in blood flow, oxygenation, and volume following extended stimulation of rodent barrel cortex. Neuroimaging 2002, 15:474-487. 49. Scheffler K, Seifritz E, Haselhorst R, Bilecen D: Titration of the BOLD effect: separation and quantification of blood volume and oxygenation changes in the human cerebral cortex during neuronal activation and ferumoxide infusion. Mag Reson Med 1999, 42:829-836. 50. Darquie A, Poline J-B, Poupon C, Saint-Jalmes H, Le Bihan D: • Transient decrease in water diffusion observed in human occipital cortex during visual stimulation. Proc Natl Acad Sci USA 2001, 98:9391-9395. Here, a change in water diffusion was observed in the human primary visual cortex during visual stimulation. This change presumably occurs due to transient swelling. It is still too early to tell whether this technique can be used for routine high-resolution fMRI studies because of low SNR. 51. Zhong J, Petroff OAC, Pleban LA, Gore JC, Prichard JW: Reversible, reproducible reduction of brain water apparent diffusion coefficient by cortical electroshocks. Mag Reson Med 1997, 37:1-6. 52. Zhong J, Kennan RP, Fulbright RK, Gore JC: Quantification of intravascular and extravascular contributions to BOLD. Magn Reson Med 1998, 40:526-536. 53. Kamei H, Iramina K, Yoshikawa K, Ueno S: Neuronal current distribution imaging using magnetic resonance. IEEE Trans Magnetics 1999, 35:4109-4111. 54. Bodurka J, Jesmanowicz A, Hyde JS, Xu H, Estkowski L, Li S-J: Current-induced magnetic resonance phase imaging. J Magn Reson 1999, 137:265-271. 55. Bodurka J, Bandettini PA: Toward direct mapping of neuronal activity: MRI detection of ultraweak, transient magnetic field changes. Magn Reson Med 2002, 47:1052-1058. 56. Pautler R, Silva A, Koretsky A: In vivo neuronal tract tracing using manganese-enhanced magnetic resonance imaging. Magn Reson Med 1998, 40:740-748. 57.

Watanabe T, Michaelis T, Frahm J: Mapping of retinal projections in the living rat using high-resolution 3D gradient-echo MRI with Mn2+-induced contrast. Mag Reson Med 2001, 46:424-429.

58. Lin CP, Tseng W, Cheng H, Chen J: Validation of diffusion tensor magnetic resonance axonal fiber imaging with registered manganese-enhanced optic tracts. Neuroimaging 2001, 14:1035-1047. 59. Lin Y-J, Koretsky AP: Manganese ion enhances T1-weighted MRI during brain activation: an approach to direct imaging of brain function. Magn Reson Med 1997, 38:378-398. 60. Duong TQ, Silva AC, Lee S-P, Kim S-G: Functional MRI of calciumdependent synaptic activity: cross correlation with CBF and BOLD measurements. Magn Reson Med 2000, 43:383-392.