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Page 37 Chapter 4 Magnetic Resonance Imaging This chapter highlights areas of interest related to magnetic resonance (MR) technology and its applications, particularly applications that involve dynamic magnetic resonance imaging (MRI). Among the hardware systems of an MRI device, the magnet, radio-frequency (RF), and gradient systems deserve particular R&D attention. For example, the development of suitable high-temperature superconductors (HTSs) and their integration into MRI magnets represents one potentially fruitful area of interaction between solid-state physicists, materials scientists, and the biomedical imaging research community. High-speed imaging pulse sequences, with particular focus on functional imaging, are discussed in this chapter, as are algorithms for image reconstruction; both are promising fields of research. Significant attention is given also to two applications for which MRI has unique potential: blood flow imaging and quantification, and functional neuroimaging based on exploiting dynamic changes in the magnetic susceptibility. Finally, this chapter highlights recent technological advances toward real-time monitoring of interventional and therapeutic procedures. MRI technology has undergone amazing strides over the last two decades, much of it due to advances from the mathematical sciences and physics. For example, Figure 4.1 demonstrates a tremendous improvement in resolution over that period. There is every reason to believe exciting progress still lies ahead.
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Page 38 Figure 4.1. Resolution, expressed as the reciprocal of imaging voxel volume, achieved in MR brain images by means of a volume head coil in scan times of less than 20 minutes and SNR sufficient to delineate anatomic structures. 4.1 Principles of Magnetic Resonance Imaging Unlike its x-ray counterparts, magnetic resonance imaging (also known as nuclear magnetic resonance (NMR) imaging) is not a transmission technique. Rather, the material imaged is itself the signal source (i.e., the macroscopic spin magnetization M from polarized water protons or other nuclei, such as 23Na or 31P). The motion of the magnetization vector of uncoupled spins, such as those for protons in water, is conveniently described in terms of the phenomenological Bloch equations: where - is the gyromagnetic ratio, H the effective field, M0 the equilibrium magnetization, and T 1 and T 2 the relaxation times. T 1 is the characteristic relaxation time for longitudinal magnetization to align with the magnetic field: following a perturbation such as a 90° RF pulse, the longitudinal magnetization typically returns to its equilibrium value, M0, with a time constant T 1. Likewise, T 2 is the characteristic time for decay of coherent magnetization in the transverse plane: the transverse magnetization decays exponentially with time constant T 2 to its equilibrium value, M° xy = 0. Both relaxation times are determined by the interaction of water or other nuclei with macromolecules in tissues. T 1 and T 2 contribute independently to the contrast between different tissues. There is, in general, no closed-form solution to equation 4.1 (although section 14.1.6 introduces two approximate solutions). Ignoring the relax-
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Page 39 Figure 4.2. Time domain nuclear magnetic resonance signal from volume element dxdydz in an object of magnetization density Mxy(t ) in the presence of a spatial encoding gradient G. ation terms (which can be done without loss of generality when formulating the imaging equations), the steady-state solution of equation 4.1 in the presence of a static polarizing field H0 = Hz corresponds to a precession of the magnetization about the field at a rate w0 = –gH0. Since the detected signal voltage is proportional to dM/dt and T1is on the order of 1 s, only components transverse to the polarizing field give rise to a detectable signal. At its core, MRI exploits the field dependence of the precession frequency by superimposing a magnetic field gradient G = (¶H0 /¶x, ¶H0 /¶ y, ¶H0 /¶ z) onto the static polarizing field H0 = H2 to spatially encode information into the signal. In this manner, the resonance frequency w becomes a function of spatial position r, according to If precession due to the polarizing field (the first term in equation 4.2) is ignored, the complex MR signal is seen to evolve as exp(-igG · rt). Following excitation by an RF pulse, and ignoring relaxation, the time-dependent signal dS(t) in a volume element dxdydz becomes (Fig. 4.2) It is convenient to express the signal as a function of the spatial frequency vector
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Page 40 where G(t') is the time-dependent spatial encoding gradient defined above, which imparts differing phase shifts to spins at different spatial locations. Equation 4.4 shows that the time variation of the spatial frequency vector is determined by the integral over the gradients; the time sequence of the RF pulses and magnetic field gradients is known as the ''pulse sequence" of the MRI acquisition. Rather than expressing the MRI signal in terms of the magnetization, which is modulated by the relaxation terms in a manner specific to the excitation scheme used, it is customary to describe the signal as a function of the density of spins in the tissue. For an object of spin density r(r) the spatial frequency signal S(k) thus is given by with integration running across the entire object. Pictorially, the spatial frequency may be regarded as the phase rotation per unit length of the object the magnetization experiences after being exposed to a gradient G(t') for some period t. One further recognizes from equation 4.5 that S(k) and r(r) are Fourier transform pairs, and thus Following an RF pulse, the transverse magnetization is subjected to the spatial encoding gradient G(t), which determines the spatial frequency vector k(t) according to equation 4.4. The manner in which k-space is scanned and the path of the k-space trajectory are determined by the waveform of the gradients G(t) and the sequence of RF pulses. In contrast with the signal used in computed tomography, the MRI signal is sampled in the spatial frequency domain, rather than in object space. The most common technique involves alternate application of two or three orthogonal gradients. Signal detection and sampling in the presence of a constant, so-called readout gradient applied along one dimension then produces a rectilinear grid of data from which the image pixel amplitudes are obtained by the discrete Fourier transform. Other imaging schemes involve radial or spiral coverage of k-space. A block diagram of a typical MRI apparatus is given in Figure 4.3. The heart of the apparatus is the magnet system, typically a superconducting
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Page 41 Figure 4.3. Simplified block diagram of a typical MRI system. solenoid. The second system is the transmit/receive assembly, typically consisting of a transmitter and power amplifier to generate the RF voltage to be fed into the transmit RF coil, thereby creating a circularly polarized RF field. The latter produces transverse magnetization, which in turn induces an RF voltage in the receive coil (which may actually serve for both transmission and reception). The ensuing signal is amplified and demodulated in the receiver electronics. Unique to the MRI device is the gradient system, which permits generation of the previously defined time-dependent gradient fields needed for spatial encoding. Both the transmit/receive and gradient systems are under the control of a data acquisition processor that ties into the main central processing unit. 4.2 Hardware 4.2.1 Magnet Systems: Current Status and Opportunities The core of an MRI apparatus is the magnet that generates the field for nuclear polarization. The magnetic field for MRI must be extremely stable (0.1 ppm per hour) and uniform (10 to 50 ppm in a sample volume of 30- to 50-cm diameter). Magnet types in current use are of the superconducting, resistive, and permanent design. The large majority of MRI units use superconducting magnets, which provide fields of high strength and stability.
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Page 42 They are, however, more expensive to manufacture and are cost-effective only for fields of 0.3 tesla (T) and higher. Most currently produced magnets are based on niobium-titanium (NbTi) alloys, which are remarkably reliable. These typically require a liquid helium cryogenic system with regular replenishment of the cryogens to keep the conductors at superconducting temperature (approximately 4.2 K). In recent years the cryogenic efficiency of such magnets has been improved so that many systems now require replenishment at 9-month intervals only, compared to monthly intervals 10 years ago, and approaches eliminating the liquid helium and using conduction cooling have been proposed. Liquid nitrogen, which previously was replenished weekly, has now been completely eliminated in some systems in favor of electrically driven refrigerators. Advances in active shielding of external magnetic fields, together with creative magnetic, mechanical, and cryogenic designs, have reduced the physical volume of these magnets by typically 40% over the past decade, and the once-delicate cryostats are now rugged enough so that mobile MRI installations can travel the world's highways without incident. Permanent and resistive magnets, both iron core and air core, are cost-effective for field strengths below approximately 0.5-0.3 T. Permanent magnets, in particular, can have very low operating costs. While permanent magnets have a weight penalty, novel designs have reduced this as well, particularly when compared to the total weight of the magnet and passive ferromagnetic shield in some other designs. Because of its ductility and ease of fabrication, the solid-state alloy of NbTi is the usual conductor used in superconductivity applications. The only other alloy used commercially is Nb3Sn, a brittle compound but one whose higher critical temperature allows it to be used in higher-field systems or at temperatures higher than 4 K. It has been used in several novel conduction-cooled open-architecture MRI magnets developed to test interventional and therapeutic applications of MRI (see Chapter 12). These systems operate at approximately 10 K. A natural next step in the evolution of MRI technology might be to wind MRI magnets with a high-temperature superconductor (HTS), one of a class of materials with critical temperatures above approximately 25 K. The first of these materials were developed by Bednorz and Muller in 1986, and ones with critical temperatures as high as 164 K have been reported. Basically ceramic alloys, these materials are constructed with powder metallurgical or thin-film techniques and are inherently brittle (similar to Nb3Sn). Considerable strides have been made in the manufacture of long production lengths of bismuth-strontium-calcium-
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Page 43 copper oxide/silver (BSCCO/Ag), which can carry higher currents than can the low-temperature superconductors at 4.2 K. In general, HTSs appear capable of carrying more current than do low-temperature superconductors at the latter's operating temperatures in fields over 14 T. A critical component of the superconducting magnet is the leads that are required to initially energize the magnet. If the magnets are persistent, these leads are retracted in order to reduce to a minimum the heat load on the system. If the magnets are conduction cooled, the power leads and any instrumentation connected to the coil are situated in the vacuum space with the windings. Because of the difficulty of disconnecting such leads in a vacuum, conduction-cooled coils usually use permanently connected leads, creating a significant heat load. There is promise that using HTS electrical leads could reduce such loads because of their natural low thermal conductivity. Another critical technology is the manufacture of true zero-resistance joints for connecting individual lengths of the magnet coils and for the persistent switch used to disconnect the magnet coils once the magnet has been energized. Creation of such components for NbTi magnets is routine but is more involved for Nb3Sn, although the use of this material in NMR spectrometers indicates that the joint technology does exist. 4.2.2 Pulsed-field MRI Systems MRI systems require a relatively high magnetic field for polarizing the magnetic moments to provide an adequate signal for signal-to-noise ratio (SNR) considerations. This polarizing operation, of itself, is relatively noncritical. Large variations, of the order of 10%, can readily be tolerated since such a variation merely shades the resultant image and can be easily compensated for if desired. However, when this same field is used for the readout operation, as is done in all existing systems, the requirements are drastically different. The field must now have homogeneities of the order of 1.0 ppm to avoid signal loss and distortion. The signal loss stems from variations in frequency within an imaging voxel1 where the resultant destructive phase interference causes enhanced signal decay, significantly reducing the sampling time available. These problems are particularly severe in high-speed systems that attempt to cover large amounts of k-space following each excitation. The limited readout time restricts the amount of coverage and the 1 Volume element, the higher-dimensional analog of the more-familiar pixel (picture element) that is the basic unit of a planar image.
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Page 44 SNR, making real-time performance impractical. In pulsed-field systems the functions of polarization and readout are separated, enabling a number of significant performance and cost advantages. A pulsed field is applied and left on for a time comparable to the relaxation time T 1, to polarize the magnetic spins. Since the magnet has considerable thermal inertia and is on only during the imaging interval, the peak field can be made relatively high, of the order of 1.0 T, without excessive temperature rise. Thus the very use of a pulsed field has a significant SNR advantage over static resistive magnets. In addition, the time of the pulse can be used for providing T 1 contrast by determining the degree of polarization buildup for the different materials. Of course 2 contrast is established in the usual way by varying the time over which the origin of k-space is scanned during the readout. As indicated, this pulsing operation is highly noncritical as to homogeneity. It therefore enables the use of low-cost simple magnets. The magnets now being used are either very large solenoids or vertical field magnets with pole pieces much larger than the imaged region. These have high stored energy and result in very limited accessibility to the patient for guided biopsies and/or therapeutic procedures. Pulsed magnets enable the use of simple configurations such as a pair of coils or a "C"-shaped core having a cross section comparable to the region of interest. This simplicity enables the construction of specialized machines for joints, head, breast, heart, and so on. The readout bias field following the pulse is made very low so that its variations will result in negligible T *2 or distortion considerations.2 Susceptibility-induced variations (which scale with readout field strength) should be negligible. Although the readout bias field is relatively noncritical, it can be a major consideration in a simple low-cost design. For example, if the readout field were 1,000 times smaller than the polarizing field, for comparable performance its homogeneity should be 1 part per 1,000 as compared to 1 ppm. Although this ratio is readily realized in a large solenoid, it remains quite difficult to achieve in a simple configuration such as a pair of coils positioned on opposite sides of the body. To achieve essentially complete immunity to inhomogeneity of the readout bias field, comparable to the immunity of the pulsed polarizing field, an oscillating bias field is used. Perhaps the biggest problem created by this approach is the radio-frequency receiving system. The desired result is to obtain an SNR commensurate 2T 2* is the effective transverse relaxation time, which determines the rate of decay of the signal in the presence of magnetic field inhomogeneity.
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Page 45 with the size of the polarizing field. This will happen only if the noise generated is due primarily to body losses, where other losses, including those of the coil, are negligible. The problem is that, at very low readout frequencies, the body losses are significantly reduced, thus making coil noise relatively more important. As a result systems using very low readout frequencies require low-loss coils. These can include coils made of cooled materials and high-temperature superconductive coils. The use of these low-loss, high-quality-factor coils introduces another problem, that of insufficient bandwidth. The small signal also results in an increased shielding problem since external interfering signals are more significant. One solution to this problem is the use of gradiometers, coil configurations that respond solely to field variations and ignore uniform fields arising from more remote sources. 4.2.3 Radio-frequency Coils for MRI Radio-frequency coils constitute the essential means of stimulating the spin systems and are the doorway for the sensing of MR data. Several research opportunities are presented by the need to improve the quality of the data and the rate of data acquisition. The data quality could be improved with better coil designs and, in some applications, cooled or superconducting coils, which result in lower noise. Better data quality could also be facilitated with multiple-acquisition coils operating in parallel. Computational Design of RF Coils Historically, RF coils have been constructed according to basic design principles learned through experience by the individual investigator. More sophisticated computational RF coil design is an area for future growth, for several reasons: · MRI is an intermediate field problem, and such fields are difficult to simulate even for quasi-static analyses. · Very high frequencies are beyond the quasi-static regime of Maxwell's equations. Wavelength effects inside the patient are important and must be considered. Designing transmitters and receivers based on an understanding of dielectric and wavelength effects, so as to enable uniform excitation and detection, is beyond the scope of quasi-static arguments. · Accurate estimates of power deposition in patients are needed for high-field, high-power imaging sequences. Since neither experimental
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Page 46 measurements nor first-order calculations can accurately predict local power deposition, accurate field calculations are required as a basis for developing safe operating guidelines. · Very high frequency coils are still tuned and matched by trial and error. As in the design of lower-frequency resonators, it would be extremely useful to have a set of independent controls for the coil's different oscillatory modes that produce circularly-polarized fields. All of these applications require the availability of a true three-dimensional solution to the time-varying fields. As yet, there has been little published on MRI applications of finite-element calculations. A few abstracts have described some small-scale problems, but none have attacked a full size problem. Generally this work has not taken a full three-dimensional approach; usually one dimension in the problem has been very coarsely sampled in the analysis. Assuming that local heating effects in a body need to be quantified, a finite-element calculation with elements on the order of 1 cm3 would be needed. A head-size coil in a shield might require something on the order of 100,000 to 200,000 elements. The scale of this problem is appropriate for even a modest modern workstation; with an iterative finite-element solver, something on the order of a week of CPU time might be required. This scale is just right for a research group to begin investigating. National supercomputer facilities exist that could greatly reduce this time. Beyond the physics of solving electromagnetic problems, issues specific to MRI complicate the interpretation of results. The best methods for visualizing three-dimensional field values are not obvious. Development of such techniques will require computational tools specific to MRI, for example, means of performing a simulated MR pulse sequence and visualizing the images created by it. Cooled Receiver Coils for MR Imaging Most SNR calculations in MRI rest on the basic assumption that the patient contributes the dominant part of the noise. This assumption is known to be incorrect in two important cases: · When coil loading increases rapidly with increasing frequency. Consequently in low-field MRI, patient contributions to noise may not dominate.
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Page 47 · When coil loading decreases rapidly with the length scale of the system. As the receiver coil is made smaller, the contribution from patient loading can be made smaller than coil contributions at any field strength. For microscopy applications the coil is the dominant source of resistance. High-temperature superconducting materials that contribute little noise when operated at liquid nitrogen temperatures deserve to be evaluated for use in MRI receiver coils. The most promising technology involves thin film HTSs. These materials can be used in large static fields provided that the plane of the thin film is parallel to the field. The limitations of thin film technology are those imposed by the vapor deposition process used to create them. Very large coils are at present not possible; the practical upper limit is about 8 cm. Alternative possibilities are tape HTSs or the use of conventional copper coils at liquid nitrogen temperature. Use of Multiple Receivers While the use of multiple acquisition coils operating in parallel could improve data quality, the technological requirements for a practical multi-coil array are formidable: beyond the issues of building a decoupled array of surface coils, there are considerable demands on the system hardware to be able to deal with data rates that are an order of magnitude higher than those from a conventional receiver coil. A flexible and powerful system architecture is needed that allows for the acquisition and manipulation of larger data streams. An active research area is the multiplexing of several RF channels through one wideband receiver, digitizer, and computer. Multiplexing is possible in both the frequency and time domains. Noise crosstalk between channels must be prevented by tight filtering near the RF coils. Crystal filters and filters built with HTS high current resonators are contenders for this task. The simplest approach for the combination of data from the array is based on a linear combination of the individual images into a composite image. Provided the receiver coil sensitivities are known and a collection of equal-noise images exists, then the least-noise combined image is a sum of images weighted by their sensitivities. However, achieving reliable and noise-free estimators of individual coil sensitivities remains a challenge that calls for more sophisticated techniques, possibly statistical ones. Alternatively, there are several adaptive techniques for image combination that act strictly on the original images without any need for filtering. Principal component analysis (PCA) can take a collection of original images
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Page 78 with some new developments that bear on future applications. Injected paramagnetic contrast agents such as gadolinium or dysprosium chelates are used with MRI to track the spatial distribution over time of the agent, from which kinetic parameters such as flow or tissue perfusion can be derived using mathematical models of various degrees of complexity. These agents effect a change in the local relaxation rate 1/T 1 that is proportional to the concentration of the contrast agent. The paramagnetic contrast agent is bound to a radionuclide tracer so that the concentration of the contrast agent in the organ or region of interest can be monitored. The sensitivity of MRI is too low for detection of the T 1 changes associated with neuroreceptortargeted ligands or antigens in general. Another class of contrast agents appropriate for MRI medical applications is the noble gases with spin 1/2, notably hyperpolarized 3He and 129Xe. As opposed to the situation with paramagnetic contrast agents, these hyperpolarized gases serve as both tracer and the source of the MRI signal. While the sensitivity of MRI is too low to detect the normal concentrations of injected nuclei, because only a few nuclei out of one million would be polarized, 10 to 20% of the nuclei are polarized with these gases, and the sensitivity increase of 10,000 to 100,000 enables their imaging. The hyperpolarized gases maintain their polarization for many minutes, or even hours if bottled in a magnetic field and stored at cryogenic temperatures. These inert gases can be used for lung imaging and, because of their finite solubility in tissue, could potentially be used for tissue perfusion quantitation as has been done with radioactive xenon, 127Xe and 133Xe. Alternatively, one can utilize a known signal decay of the nuclear spin relaxation rate, 1/T1, to determine flow at equilibrium during constant infusion, probably through inhalation of a dilute mixture of the hyperpolarized gases. 4.4.6 Microscopic Imaging There is no well-defined boundary between standard macroscopic imaging and microscopic imaging. A common working definition places that boundary at the best resolution attainable with the usual clinical MRI systems, perhaps about 300 mm for protons. The attainable resolution using surface coils can be less than 50 /m, so that MRI covers a volume range of about 106, with a corresponding range of total numbers of nuclei and hence signal strength. The potential applications range from the improved delineation of anatomical structure already resolvable by MRI to new classes of features in tissues. Examples of specialized applications that may become important include visualization of cortical layers, the precise boundaries of functional regions in the brain, and microangiography at the level of arterioles and
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Page 79 venules. All of these applications may be augmented, or made possible, by the use of magnetic contrast agents to allow the identification of specific objects, such as blood vessels and tissues of particular types. Under various circumstances, resolution may be limited by the achievable SNR, digitization (during acquisition or processing), bulk motion, or molecular diffusion, as well as by other problems and artifacts typically encountered in MRI. Resolution It is common in practice to confound true image resolution with digital resolution (the dimensions of a pixel or voxel) and to emphasize pixel resolution in a thick slice. These practical working definitions can give rise to disagreements when more rigorous analyses of instrument performance and image interpretation are carried out. There is no fundamental physical limit to resolution in MRI. The limits encountered in practice are complicated functions of the features of the apparatus, acquisition and processing methods, and object characteristics. In a given experiment, the limits may be set by the SNR in a single voxel, by digital resolution, by intrinsic line widths (homogeneous or heterogeneous broadening, by T 2 or susceptibility effects, for example), by diffusional effects, or by local or bulk motions. Signal-to-Noise Ratios SNR may be increased by using higher magnetic fields, improved RF coil designs (including cooled coils), smaller (external, intrusive, or implanted) coils, more efficient acquisition methods, and special processing methods. When the contrast-to-noise ratio or the T 1/T 2 ratio is important, contrast agents may increase the effectiveness (the contrast attainable in a fixed time). Small objects can now be imaged with voxel resolution in the range from 100 mm3 to 1000 mm3, using receiver coils of the order of millimeters in diameter at magnetic fields of several tesla. It can be anticipated that regions near surfaces or implanted coils of similar dimensions will be imaged in vitro at similar resolution (about 10 ym isotropically) in the near future. Ultramicro coils (10- to 100-mm diameter, with integral preamplifier) may improve the volume resolution by about an order of magnitude because they increase the SNR by a similar factor. Because in the small-coil limit the noise arises predominantly from the coil rather than the sample, the recent development of superconducting receiver coils has afforded dramatic enhancements in sensitivity that, in turn, can be traded for improved resolution by decreasing voxel size. Gradients As the SNR is increased, the gradient strengths required to encode image information become a problem. High-amplitude gradient coils to
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Page 80 surround small objects are readily constructed and operated, but in vivo applications require special structures, efficient encoding methods to minimize the gradient amplitudes required, and low-duty cycles or effective cooling to avoid local heating. These considerations probably limit resolution to the order of a few tenths of a millimeter for the immediate future in human application. Diffusion Molecular diffusion is becoming a useful source of contrast and microstructural information in studies at ordinary resolution. It plays the same role in microscopy, although unwanted diffusion damping may be a more important factor in microscopy because of the large gradients. As resolution of tens of micrometers is approached, diffusion displacements of nuclei become more important, and below a 10-mm pixel size molecular diffusion becomes the dominant consideration in experimental design. Motion The effects of motion may be counteracted by faster image acquisition, gated acquisition (to the cardiac cycle, for example), and a variety of other tracking and processing methods already used in clinical practice and research. Postprocessing, such as realignment of images to compensate for rigid body motions, is becoming more practical as readily available computing power increases. Newer methods, such as (k, t)-space imaging, may provide a more general solution to the problem. Within some regions of the brain, motion amplitudes may be only a few hundred micrometers or less. Elsewhere in the body the problems will almost always be greater, and specific solutions may be required. It should be noted that the coordinate system is defined by the magnetic field gradients, not by spatial location itself. Movement of gradient coils can be mistaken for object motion, and gradients that track object motion would eliminate primary motion artifacts, but are not now practical. Future Applications of in vivo MRI Microscopy A rich variety of structural and functional features becomes visible below the millimeter level. Their visualization will usually be possible with apparatus and techniques developed for specific structures in specific locations, and for studies of a well-defined problem of function or diagnosis. 4.4.7 Research Opportunities Related to Applying Dynamic MRI Blood Flow All MR velocity studies will benefit from hardware improve- ments that affect flow image quality. These include dedicated RF coils,
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Page 81 high-performance magnetic field gradients, reduced eddy current effects, and improved spatial homogeneity, to name a few. Software improvements to enable spatially selective RF pulses are also vital. Postprocessing and display of velocity information are currently tedious and impractical. Improvements in this area must also permit the exploration of additional information in the data, such as the extraction of pressure values and shear/stress forces. Some specific research opportunities include the following: · Development of techniques for rapid measurement of instantaneous velocity in three-dimensional space (six-dimensional problem); · Modeling of complex flow and its implications for the vascular MR signal; and · Development of methods for extracting parameters of physiologic relevance, including shear stress, distensibility, and turbulence. Diffusion Imaging One field of research regards the implementation of diffusion tensor imaging (DTI) with EPI on conventional systems for clinical use. The implementation of EPI would require improvements in the gradient hardware (eddy current compensation and high-amplitude, high-slew-rate gradients), which would also benefit diffusion accuracy. Further, understanding the mechanisms underlying the diffusion values in tissues would be highly desirable. Most diffusion values are about one order of magnitude smaller in tissues than in pure water. Part of this difference can be explained by the tissue microstructure, in terms of obstacles, fibers, or membranes, but some of the differences may be artifacts. Matching MRI diffusion measurements by DTI with direct measurements using microelectrodes in tissues or animal preparations would be extremely useful. With those techniques the existence of restricted diffusion effects, requiring ultrashort diffusion times, could be demonstrated. Also the mechanisms of anisotropy in brain white matter could be better understood, as well as those involved in acute brain ischemia. The clinical value of diffusion MRI in stroke needs more evaluation. What is the prognostic value of diffusion in terms of patient recovery? How would repeated diffusion measurements help monitor the effects of drug therapy? On the other hand, would anisotropic diffusion studies benefit management of patients with myelin disorders (e.g., brain development retardation, multiple sclerosis)? More generally, what would be the role of DTI in the evaluation of brain diseases? Still to be done are in vivo studies demonstrating
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Page 82 the feasibility of the application of DTI to temperature imaging in interventional MRI, as well as evaluations of the effects of blood flow and tissue denaturation on the diffusion measurements. Some specific research opportunities related to diffusion imaging include: · Establishment and experimental validation of a molecular model for anisotropic diffusion in tissues; · Development of improved approaches for spatially localized measurement of diffusion coefficients in vivo; · Modeling of heat dissipation in tissue; and · In vivo measurement of tissue fiber orientation. Other Tissue Parameters · Development of accurate measurement techniques for quantitative relaxation times and their interpretation in terms of clinical diagnosis. · Development of sophisticated segmentation techniques based on multiple parametric acquisitions. · Extraction and meaningful display of strain maps of cardiac function. Functional Brain MRI · Development of methods for monitoring data quality during a scanning session and for ensuring that functional activation is being observed. · Development of new experimental protocols, especially for use with complex stimuli (e.g., visual presentations). · Modeling and experimental verification of the biophysical contrast-tonoise mechanisms induced by neuronal activation and their dependence on magnetic field strength. · Establishment of detailed biophysical models to understand the stimulus response permitting separation of vascular from parenchymal changes and arterial from venous changes, taking into account parameters such as blood volume, blood flow, and so on. · Evaluation of the nature of various sources of noise (e.g., stochastic, physiological, instrument instability) and development of strategies for their minimization.
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Page 83 Multinuclear MRI · Evaluation of polarized noble gases as tracers of pulmonary function and tissue perfusion (e.g., in muscle and in the brain and other organs). Microscopic Imaging · Exploration of the theoretical limit of spatial resolution and its dependence on key parameters, including diffusion and detection sensitivity. · Development of improved means for monitoring of and correcting for the effects of motion, which currently limits resolution of in vivo MR microscopy. 4.4.8 Suggested Reading on Applications of Dynamic MRI Blood Flow 1. Bryant, D.J., Payne, J.A., Firmin, D.N., et al., Measurement of flow with NMR imaging using gradient pulse and phase difference technique, J. Comput. Asst. Tomog. 8 (1984), 588-593. 2. Caro, C.G., Pedley, T.J., Schroter, R.C., and Seed, W.A., The Mechanics of Circulation, Oxford University Press, Oxford, 1978. 3. van Dijk, P., Direct cardiac NMR imaging of heart wall and blood flow velocity, J. Comput. Asst. Tomog. 8 (1984), 429-436. 4. Dumoulin, C.L., Souza, S.P., Hardy, C.J., and Ash, S.A., Quantitative measurement of blood flow using cylindrically localized Fourier velocity encoding, Magn. Reson. Med. 21 (1991), 242-250. 5. Dumoulin, C.L., Souza, S.P., Walker, M.F., and Wagle, W., Three-dimensional phase contrast angiography, Magn. Reson. Med. 9 (1989), 139-149. 6. Feinberg, D.A., Crooks, L.E., Sheldon, P., Hoenninger III, J., Watts, J., and Arakawa, M., Magnetic resonance imaging the velocity vector components of fluid flow, Magn. Reson. Med. 2 (1985), 555-566. 7. Mohiaddin, R.H., Firmin, D.N., Underwood, S.R., et al., Aortic flow wave velocity: The effect of age and disease, Magn. Reson. Imaging 7 (suppl. 1) (1989), 119.
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Representative terms from entire chapter: