Functional Imaging of Cognition
Functional Imaging Reflects the Metabolic Demand of Neural Activity
Functional Imaging Is Used to Probe Cognitive Processes
Functional Imaging Has Limitations
THE ABILITY OF NEURO-IMAGING to observe areas of the human brain that are active during cognitive processes has helped to stimulate the current interest in the biological underpinnings of cognitive functioning. Because invasive experiments cannot be done ethically on humans, research on the biological basis of cognitive function was until quite recently confined to laboratory animals and clinical studies of patients with cognitive disorders.
The development of techniques such as functional magnetic resonance imaging (fMRI) has made it possible to study human subjects, affording unprecedented views of the complexities of the intact working brain. Imaging of the living brain allows us to explore the behavioral significance of local neural circuits, such as cortical columns, as well as observe large-scale systems of interconnected brain regions concerned with specific mental processes such as seeing, hearing, feeling, moving, talking, and thinking.
Functional Imaging Reflects the Metabolic Demand of Neural Activity
Functional Imaging Emerged from Studies of Blood Flow
Functional imaging evolved out of seminal studies in the late 1940s by Seymour Kety and Carl F. Schmidt, who succeeded in measuring blood flow in the living brain. Although Charles S. Roy and Charles S. Sherrington earlier had found a relationship between blood flow and brain metabolism, Kety and Schmidt were the first to quantify cerebral blood flow noninvasively.
To accomplish this task, Kety and Schmidt measured the rate of cerebral blood flow by having subjects inhale nitrous oxide, a metabolically inert gas, and measuring its outflow concentration from the jugular vein (Box 20–1). In a series of landmark studies they evaluated how blood flow from the intact brain varied in different metabolic states, such as sleep and wakefulness, and in so doing they laid the foundations for modern functional imaging.
These early experiments measured only the total level of activity of the entire brain, however. They could not provide information about which parts of the brain were active, nor could they tell us whether some brain areas became more active while others became less active under set conditions.
A significant advance came in the 1970s with the introduction of positron emission tomography (PET) by Michel Ter-Pogossian, Michael Phelps, and Louis Sokoloff (Box 20–2). In the 1980s Marcus Raichle collaborated with Michael Posner to visualize the brain activity of subjects engaged in complex tasks of thought and language, thereby demonstrating that PET can be used to explore cognitive functioning.
Box 20-1 Application of the Fick Principle to Brain Metabolism
Devised as a technique for measuring cardiac output by Adolf Eugen Fick, the Fick principle states that an organ must receive blood at a rate that is equal to the rate at which the organ metabolizes a constituent of blood, divided by the concentration of that constituent.
The essence of the Fick principle is that blood flow to an organ can be calculated using a marker substance. The principle may be applied in many ways. For example, if blood flow to an organ is known, together with the arterial and venous concentrations of the marker substance, then the uptake or metabolism by the organ may be calculated.
Seymour Kety and Carl F. Schmidt adapted the Fick principle so that it could be applied to the brain and showed that it can be used to measure cerebral blood flow. The Fick Principle has also been used to explain BOLD (Blood Oxygen Level Dependent) fMRI. BOLD fMRI detects changes in deoxyhemoglobin content within a unit volume of brain. As can be derived from the Fick principle, deoxyhemoglobin concentration is proportional to the cerebral metabolic rate of oxygen (CMRO2) divided by cerebral blood flow (CBF).
A further advance in functional imaging occurred in 1990 when Seiji Ogawa and David Tank discovered that magnetic resonance imaging (MRI) can be made sensitive to changes in deoxyhemoglobin that are caused when neurons change their metabolic rates. They exploited the fact, first discovered in 1936 by Linus Pauling, that when oxyhemoglobin is converted to deoxyhemoglobin (by stripping hemoglobin of its four oxygen molecules), it becomes paramagnetic.
In particular, they showed that in MRI images of the hippocampal formation of anesthetized rodents areas of increased vascularity appeared darker than areas with less vascularity. When rodents breathed 100% oxygen, the image intensities in the hippocampus were brighter, thereby suggesting that these differences in intensity were caused by changes in blood oxygenation. Finally, Ogawa went on to link these differences in image intensity with metabolism. He found that systematic pharmacological alteration of basal brain metabolism in anesthetized animals induced a corresponding increase in the image intensity (Box 20–3).
In brain regions with increased metabolism the flow of oxygenated blood is greater than the consumption of oxygen, and thus leads to a relative decrease in deoxyhemoglobin. MRI areas with increased metabolism and flow of oxygenated blood appear brighter than regions that are not experiencing increased metabolism (Box 20–3). This form of functional MRI has been termed BOLD (blood oxygen level dependent) imaging.
Positron emission tomography (PET) imaging requires the introduction into the brain of substances tagged with radionuclides that emit positrons (positively charged electrons). Commonly used substances are 11C, 18F, 15O, and 13N. The synthesis of compounds with these radionuclides does not result in the loss of biological activity; thus H215O behaves like H216O and 18F-deoxyglucose like deoxyglucose.
The radionuclides are produced in a cyclotron, which adds protons into the nuclei of atoms. For example, bombarding oxygen with hydrogen ions makes 18F. Incorporation of an extra proton into the nucleus produces an unstable nucleus.
These unstable radionuclides can be detected when the extra proton spontaneously breaks down into two particles: (1) a neutron, which remains within the nucleus because a stable nucleus can contain extra neutrons, and (2) a positron, a particle that travels away from the nucleus at the speed of light, dissipating energy as it goes. The positron eventually collides with an electron, and the collision leads to their mutual annihilation and the emission of two gamma rays (high-energy photons) in opposite directions (Figure 20-2A).
Figure 20-2A Emission of gamma rays. The nucleus of an unstable radionuclide emits a positron. The positron travels a certain distance before it collides with an electron and is annihilated, emitting two gamma rays that travel in precisely opposite directions. The site of positron annihilation that is imaged may be a few millimeters from the site of origin. For example, the average distance between the site of origin and annihilation is 2 mm for 18F and 3 mm for 15O. The distance between the emitting nucleus and the site where the positron is annihilated is an absolute limit on the spatial resolution of PET scan images. (Adapted, with permission, from Oldendorf 1980.)
PET scanners contain arrays of gamma ray detectors (scintillation crystals coupled to photomultiplier tubes) encircling the subject’s head (Figure 20-2B). The two gamma rays emitted by the annihilation of a positron and electron ultimately reach pairs of coincidence detectors that record an event when, and only when, two gamma rays are detected simultaneously.
Figure 20-2C PET image. PET produces an image showing the areas of heightened neural activity as revealed by the radionuclides.
A coincident pair of gamma ray emissions is detected along a line in one plane or slice. The site where the positron is annihilated is the site detected by the scanner. Multiple positron-electron annihilations are pinpointed by monitoring coincident gamma rays in multiple slices. Clusters of annihilations indicate increased neural activity, which is mapped onto the brain in the final PET image (Figure 20-2C).
The distance between the site of annihilation and the emitting nucleus, which can be several millimeters, limits the spatial resolution of the method, which is typically 6 to 8 mm. The temporal resolution of PET imaging is limited by the rate at which positrons are emitted, which ranges from minutes to hours depending on the radionuclide used and the compound in which it is incorporated.
Functional Imaging Reflects Energy Metabolism
As these discussions make clear, functional imaging does not measure neural activity but rather reflects energy metabolism, best defined as the rate at which mitochondria produce adenosine triphosphate (ATP). Because direct imaging of ATP production is difficult, functional imaging assesses correlates of energy metabolism that can be visualized with clinical imaging devices (Figure 20-1).
Figure 20-1 Relationship to brain metabolism. The energy metabolism of neurons is influenced by changes in synaptic activity or synaptic strength. Shifts in metabolism are associated with local increases in cerebral blood flow, glucose uptake, and cerebral blood volume, and a decrease in deoxyhemoglobin content. These different changes are detected with different techniques. Imaging techniques: fMRI, functional magnetic resonance imaging; PET, positron emission tomography; SPECT, single-photon emission computed tomography.
A surprisingly large amount of a neuron’s total energy metabolism, approximately one-half, is devoted simply to maintaining the resting membrane potential—the electric potential across the cell membrane. Therefore any shift in the membrane potential will affect the rate of energy metabolism and influence functional imaging measures. The membrane potential changes when a cell fires an action potential and also in response to subthreshold excitatory or inhibitory synaptic potentials.
The remaining half of a neuron’s energy metabolism is devoted to other biochemical processes, and alterations in these pathways also affect functional imaging measures, although typically on a slower time scale. These biochemical processes include all of the molecular reactions required for normal synaptic function: vesicle recycling, recruitment of second-messenger cascades, local protein synthesis, axonal transport, and transmitter release. Thus functional imaging in principle can measure the transient effect an external stimulus has on the electrical activity of neurons, as well as the more permanent effect of a disease process on neuronal biochemistry.
The development of functional magnetic resonance imaging (fMRI) emerged from a chain of discoveries that began in 1937 with the description of molecular beam magnetic resonance by Isidor Rabi and the discovery in 1945 of nuclear magnetic resonance (NMR), made independently by Edward Purcell and Felix Bloch. In 1949 Erwin Hahn observed that NMR decays differentially depending on the chemical makeup of an object, the key phenomenon that has made fMRI possible.
MRI scanners consist of several components. The first component is a superconducting magnet that provides a powerful and very uniform magnetic field (1.5 tesla for a standard clinical MRI scanner). Each water proton in the body rotates around its axis and acts like a small bar magnet. Water protons normally have random directions so the tissue essentially has no net magnetization. However, when placed in a magnetic field the protons become aligned (Figure 20-3A).
Figure 20-3 (Opposite) Magnetic resonance imaging.
A. Water protons spin around their axes, creating individual magnetic fields with random directions (1). When a vertical magnetic field is applied to the tissue, the protons align with it to create a net magnetic field that is also vertical but very small and difficult to detect (2). A radio frequency pulse applied in a second (horizontal) direction makes the protons wobble, or precess, around their vertical axes (3). Summed across all of the individual water protons, this creates a net magnetic field that changes in time and gives rise to an electric current that is ultimately measured in MRI (4).
B. An MRI measurement begins by placing the subject in a vertical magnetic field. With the protons aligned vertically, a horizontal radio frequency pulse is applied to tip the protons so that they rotate in the horizontal plane synchronously, or “in phase” with one another (1). The horizontal pulse is then turned off (2), and the rotating protons begin to move out of phase with one another—they “dephase.” Dephasing occurs relatively quickly and leads to a decrease or decay in the measured current. The time constant of this decay is called T2* (approximately 30 ms). After withdrawal of the horizontal pulse the protons realign with the vertical magnetic field (3–5). This “righting” or recovery of the vertical magnetization occurs more slowly than the dephasing. The time constant of the recovery is called T1 (several seconds). The entire process can be repeated many times to yield a time series of measurements that reflect changes in the rates of decay and recovery.
The second component is a radio frequency coil (or RF coil), a specially designed coil of wire placed near the subject. A brief, rapidly alternating electrical current in the RF coil generates a rapidly varying magnetic field because of Ampere’s law. This second magnetic field is superimposed with the scanner’s main magnetic field. The alternating electrical current in the RF coil is called a radio frequency pulse (or RF pulse) because it alternates at a frequency comparable to FM radio frequencies.
The magnetic field induced by the RF pulse causes the protons to start wobbling around their axes (Figure 20-3A), much as a spinning top wobbles around its axis when the force of gravity competes with its spin. This wobbling is called precession. The protons continue to precess after the RF pulse has been turned off.
Summed across all of the individual water protons, the precession creates a rotating magnetic field that changes in time (Figure 20-3A) and, according to Faraday’s law, generates an alternating electric current back in the RF coil. It is this electric current that is measured in MRI (Figure 20-3B).
The amplitude of the measured electric current decays over time at a rate that is dependent upon a number of factors, including the type of tissue in which the protons are embedded. Thus, differences in tissue type appear as different intensities in the resulting images.
The third component of an MRI scanner is the magnetic gradient coils. One of the most important developments in MRI is the ability to make three-dimensional images of the body. This is accomplished by using magnetic gradients, magnetic fields in which the strength of the field changes gradually along an axis.
It is beyond the scope of this chapter to explain in detail how two-dimensional images (or three-dimensional volumes) are acquired with MRI, but the basic idea is that controlling the magnetic gradients allows one to measure the MRI signal (the electric current in the RF coil) at a large number of adjacent locations, each corresponding to a small volume (or voxel) of tissue.
Functional MRI primarily measures changes in the relative amount of deoxyhemoglobin within each voxel (Figure 20-4). When neurons are active, the supply of oxygenated blood to the active region increases. For reasons that are still unclear, the delivery of oxygenated hemoglobin is greater than local oxygen consumption, resulting in a greater proportion of oxygenated to deoxygenated hemoglobin.
Oxygenated and deoxygenated hemoglobin have different magnetic properties. Hemoglobin contains iron, which is exposed when oxygen is stripped from the hemoglobin molecule. The presence of deoxyhemoglobin introduces an inhomogeneity in the nearby magnetic field. Some water protons (those that are near a deoxyhemoglobin molecule) now experience a magnetic field strength that is slightly different from the other water protons.
Greater inhomogeneity causes the protons to desynchronize more rapidly resulting in a more rapid decay time (T 2*). When there is an increase in oxygenated blood in areas with greater neuronal activity, and hence a more homogeneous magnetic field, the result is a longer T 2* decay time, and brighter image intensity.
Like PET scanning, fMRI is sensitive to the increased blood flow associated with neural activity. This technique has several advantages over PET scanning, however. It requires no injection of foreign substances into the bloodstream (fMRI uses endogenous hemoglobin as its marker). It also offers finer spatial and temporal resolution than PET.
For example, fMRI has been used to visualize the ocular dominance columns in human V1, which requires a spatial resolution of less than a millimeter, and it has been used to estimate differences in the timing of neural activity with a temporal resolution of approximately 100 ms. Sub-millimeter and sub-millisecond resolutions are not yet routine practice but have been demonstrated convincingly. The fMRI image in Figure 20-4 was acquired with a spatial resolution of 1 mm.
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