Magnetic Resonance Imaging: Principles and Techniques: Lessons for Clinicians

The development of magnetic resonance imaging (MRI) for use in medical investigation has provided a huge forward leap in the field of diagnosis, particularly with avoidance of exposure to potentially dangerous ionizing radiation. With decreasing costs and better availability, the use of MRI is becoming ever more pervasive throughout clinical practice. Understanding the principles underlying this imaging modality and its multiple applications can be used to appreciate the benefits and limitations of its use, further informing clinical decision-making.In this article, the principles of MRI are reviewed, with further discussion of specific clinical applications such as parallel, diffusion-weighted, and magnetization transfer imaging. MR spectroscopy is also considered, with an overview of key metabolites and how they may be interpreted. Finally, a brief view on how the use of MRI will change over the coming years is presented. The development of magnetic resonance imaging (MRI) for use in medical investigation has provided a huge forward leap in the field of diagnosis, particularly with avoidance of exposure to potentially dangerous ionizing radiation. With decreasing costs and better availability, the use of MRI is becoming ever more pervasive throughout clinical practice. Understanding the principles underlying this imaging modality and its multiple applications can be used to appreciate the benefits and limitations of its use, further informing clinical decision-making. In this article, the principles of MRI are reviewed, with further discussion of specific clinical applications such as parallel, diffusion-weighted, and magnetization transfer imaging. MR spectroscopy is also considered, with an overview of key metabolites and how they may be interpreted. Finally, a brief view on how the use of MRI will change over the coming years is presented. The nuclear magnetic resonance (NMR) phenomenon was first described experimentally by both Bloch and Purcell in 1946, for which they were both awarded the Nobel Prize for Physics in 1952.1Bloch F. Hansen W.W. Packard M.E. Nuclear induction.Phys Rev. 1946; 69: 127Crossref Scopus (579) Google Scholar, 2Purcell E.M. Torrey H.C. Pound R.V. Resonance absorption by nuclear magnetic moments in a solid.Phys Rev. 1946; 69: 37-38Crossref Scopus (1985) Google Scholar The technique has rapidly evolved since then, following the introduction of wide-bore superconducting magnets (approximately 30 years ago), allowing development of clinical applications. The first clinical magnetic resonance images were produced in Nottingham and Aberdeen in 1980, and magnetic resonance imaging (MRI) is now a widely available, powerful clinical tool.3Hawkes R.C. Holland G.N. Moore W.S. Worthington B.S. Nuclear magnetic resonance (NMR) tomography of the brain: a preliminary clinical assessment with demonstration of pathology.J Comput Assist Tomogr. 1980; 4: 577-586Crossref PubMed Scopus (127) Google Scholar, 4Smith F.W. Hutchison J.M. Mallard J.R. et al.Oesophageal carcinoma demonstrated by whole-body nuclear magnetic resonance imaging.Br Med J (Clin Res Ed). 1981; 282: 510-512Crossref PubMed Scopus (39) Google Scholar This article covers a brief synopsis of basic principles in MRI, followed by an overview of current applications in medical practice. All atomic nuclei consist of protons and neutrons, with a net positive charge. Certain atomic nuclei, such as the hydrogen nucleus, 1H, or the phosphorus nucleus, 31P, possess a property known as “spin”, dependent on the number of protons. This can be conceived as the nucleus spinning around its own axis although this is a mathematical analogy. The nucleus itself does not spin in the classical meaning but by virtue of its constituent parts induces a magnetic moment, generating a local magnetic field with north and south poles. The quantum mechanical description of this dipolar magnet is analogous to classical mechanics of spinning objects. The dipole itself is analogous to a bar magnet, with magnetic poles aligning along its axis of rotation (Figure 1).5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar Application of a strong, external magnetic field (B0) aligns the nucleus either in parallel with or perpendicular to the external field. A liquid solution containing many nuclear spins, placed within the B0 field, will contain nuclear spins in one of two energy states: a low-energy state (oriented parallel to the magnetic field) or a high-energy state (orientated perpendicular to the magnetic field direction). In solids or liquids, there would tend to be an excess of spins in the same direction as B0. Although a bar magnet would orientate completely parallel or antiparallel to the field, the nucleus has an angular momentum due to its rotation, so it will rotate, or precess, around the B0 axis (Figure 2). This behavior is often compared to the wobbling motion of a gyroscope under the influence of the Earth's magnetic field and explains the use of “spin” to explain what is in reality a quantum mechanical phenomenon. The velocity of rotation around the field direction is the Larmor frequency. This is proportional to the field strength, and is described by the Larmor equation (Figure 3).5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google ScholarFigure 3Larmor equation. ω0 = angular frequency of the protons, γ is the gyromagnetic ratio, a constant fixed for a specific nucleus and B0 is the field strength. ΔE = γhB0/2πView Large Image Figure ViewerDownload (PPT) Nuclei that possess spin can be excited within the static magnetic field, B0, by application of a second radiofrequency (RF) magnetic field B1, applied perpendicular to B0. The RF energy is usually applied in short pulses, each lasting microseconds. The absorption of energy by the nucleus causes a transition from higher to lower energy levels and vice versa on relaxation. The energy absorbed (and subsequently emitted) by the nuclei induces a voltage that can be detected by a suitably tuned coil of wire, amplified and displayed as the “free-induction decay” (FID). In the absence of continued RF pulsation, relaxation processes will return the system to thermal equilibrium. Therefore, each nucleus will resonate at a characteristic frequency when placed within the same magnetic field.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar The energy required to induce transition between energy levels is the energy difference between the two nuclear spin states. This depends on the strength of the B0 magnetic field the nuclei are subjected to (Figure 4). Application of an RF pulse at the resonant frequency generates a FID. In practice, multiple RF pulses are applied to obtain multiple FIDs, which are then averaged to improve the signal-to-noise ratio (SNR). The signal-averaged FID is a time-domain signal. It will be made up of contributions from different nuclei within the environment being studied (e.g. free water and 1H bound to tissue). The signal-averaged FID can be resolved by a mathematical process known as Fourier transformation, into either an image (MRI) or a frequency spectrum, providing biochemical information (Figure 5).5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google ScholarFigure 5The free induction decay (FID) and Fourier transformation to generate MR images or MR spectra.View Large Image Figure ViewerDownload (PPT) Localizing the MR signal spatially to a region of interest requires the use of gradients. These are additional spatially linear variations in the static field strength. Gradients can be applied in any orthogonal direction using the three sets of gradient coils, Gx, Gy, and Gz, within the MR system. Faster or slower precession is detected as higher or lower MR signal. Thus, the frequency measurements can be used to distinguish MR signals at different positions in space and enable image reconstruction in three dimensions (Figure 6).5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar The transmitter and receiver coils may be either separate or individual pieces of hardware, depending on the area of body under examination and the experiment being performed. The applied B1 pulse is applied by an enveloping transmitter coil, which uniformly surrounds the area of interest, such as a head coil. The receiver coil consists of a loop of wire, which may either be placed directly over the region of interest or combined within the transmitter coil. Phased-array coils involve a number of coils receiving MR signal simultaneously and independently from a single excitation. If each coil is connected to a separate receiver, then the noise between the coils is uncorrelated, resulting in a higher signal-to-noise ratio than if the coils were just connected to one receiver. Mathematical algorithms can then be employed to combine the data from the individual coils to generate an optimum reconstructed Image.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar Parallel imaging is an MR technique designed to reduce scan time. Sensitivity encoding (SENSE™, Philips) and simultaneous acquisition of spatial harmonics (SMASH) are two such examples. SENSE works through under-sampling of the MR data and by collecting data simultaneously from multiple imaging coils. Reconstruction of the data requires an accurate knowledge of the individual coil sensitivities prior to the acquisition of the data. Therefore, a reference scan acquiring low resolution individual coil data is acquired prior to the main imaging sequence. Thus, a SENSE factor of 2 may reduce imaging time by up to 50%. However, with higher SENSE factors there may be a diminishing amount of MR signal that is recorded.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar Current diagnostic MRI scanners use cryogenic superconducting magnets in the range of 0.5 Tesla (T) to 1.5 T. By comparison, the Earth's magnetic field is 0.5 Gauss (G), which is equivalent to 0.00005 T. Cooling the magnet to a temperature close to absolute zero (0 K) allows such huge currents to be conducted; this is most commonly performed via immersion in liquid helium. Until recently, most clinical research was conducted at a field strength of 1.5 T. However, 3 T systems are now widely available and are being used regularly in the research setting, where the capabilities of 3 T systems are being explored and optimized. The advantages of higher field strength systems include improved signal-to-noise ratio (SNR), higher spectral, spatial, and temporal resolution, and improved quantification. The improved SNR can be traded to allow a reduced imaging time. Inherent disadvantages include magnetic susceptibility, eddy current artifacts, and magnetic field instability.6Di Costanzo A. Trojsi F. Tosetti M. et al.High-field proton MRS of human brain.Eur J Radiol. 2003; 48: 146-153Abstract Full Text Full Text PDF PubMed Scopus (65) Google Scholar, 7Soher B.J. Dale B.M. Merkle E.M. A review of MR physics: 3 T versus 1.5 T.Magn Reson Imaging Clin N Am. 2007; 15: 277-290Abstract Full Text Full Text PDF PubMed Scopus (168) Google Scholar Magnetic susceptibility is the degree of magnetization that a tissue or material exhibits in response to a magnetic field. This may have either a beneficial or deleterious effect on the overall image quality. Magnetic susceptibility artifacts are more prominent at 3 T compared to 1.5 T. The phenomenon may be beneficial in functional or diffusion MRI by improving tissue contrasts, but disadvantageous by producing signal voids at air/tissue interfaces in diffusion sequences. An eddy current is an induced current generated due to the interaction between the rapidly changing magnet field and the conducting structures within the MRI scanner. Eddy currents may lead to perturbations in the gradient field, reducing resolution of the subsequent MR Image.7Soher B.J. Dale B.M. Merkle E.M. A review of MR physics: 3 T versus 1.5 T.Magn Reson Imaging Clin N Am. 2007; 15: 277-290Abstract Full Text Full Text PDF PubMed Scopus (168) Google Scholar Relaxation is the term used to describe the process by which a nuclear “spin” returns to thermal equilibrium after absorbing RF energy. There are two types of relaxation, longitudinal and transverse relaxations, and these are described by the time constants, T1 and T2, respectively.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar T1 is also known as “spin-lattice relaxation”, whereby the “lattice” is the surrounding nucleus environment. As longitudinal relaxation occurs, energy is dissipated into the lattice. T1 is the length of time taken for the system to return 63% toward thermal equilibrium following an RF pulse as an exponential function of time. T1 can be manipulated by varying the times between RF pulses, the repetition time (TR). Water and cerebrospinal fluid (CSF) have long T1 values (3000–5000 ms), and thus they appear dark on T1-weighted images, while fat has a short T1 value (260 ms) and appears bright on T1-weighted images.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar Relaxation processes may also redistribute energy among the nuclei within a spin system, without the whole spin system losing energy. Thus, when a RF pulse is applied, nuclei align predominantly along the axis of the applied energy. On relaxation, there is dephasing of nuclei orientations as energy is transferred between the nuclei and there is reduction in the resultant field direction, with a more random arrangement of alignments. This is T2, termed transverse relaxation, because it is a measure of how fast the spins exchange energy in the “xy” plane. T2 is also known as “spin-spin” relaxation.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar MT indirectly allows measurement of bound and free water compartments in the brain. It can be affected by variations in membrane fluidity, heavy metal concentration, and total water content.8Rovira A. Cordoba J. Sanpedro F. Grive E. Rovira-Gols A. Alonso J. Normalization of T2 signal abnormalities in hemispheric white matter with liver transplant.Neurology. 2002; 59: 335-341Crossref PubMed Scopus (69) Google Scholar, 9Rovira A. Grive E. Pedraza S. Rovira A. Alonso J. Magnetization transfer ratio values and proton MR spectroscopy of normal-appearing cerebral white matter in patients with liver cirrhosis.Am J Neuroradiol. 2001; 22: 1137-1142PubMed Google Scholar MT itself is a technique for manipulating tissue contrast.10Hajnal J.V. Baudouin C.J. Oatridge A. Young I.R. Bydder G.M. Desig and implementation of magnetization transfer pulse sequences for clinical use.J Comput Assist Tomogr. 1992; 16: 7-18Crossref PubMed Scopus (131) Google Scholar, 11Wolff S.D. Balaban R.S. Magnetization transfer contrast (MTC) and tissue water proton relaxation in vivo.Magn Reson Med. 1989; 10: 135-144Crossref PubMed Scopus (1230) Google Scholar In addition to enabling acquisition of images with enhanced contrast, techniques employing MT allow measurement of MT ratios (MTR) (Figure 7). MTR is a quantitative tissue characteristic reflecting the behavior of normally MR-invisible protons bound to macromolecules. MTR measurement can detect parenchymal changes in the brain that may not be seen using standard MR techniques.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google Scholar In essence, protons in tissues exist in two pools, free and bound. Mobile protons, such as those found in body water make up the free pool; it has a narrow spectral line with relatively long T1 and T2 relaxation times (Figure 8). The majority of signal in conventional MR applications comes from the free pool, as the range for MR excitation frequency is narrow and centered on these mobile protons. A second pool of protons bound in proteins and other macromolecules or membranes is referred to as being MR invisible, as it is not typically within the excitation frequency range used. This pool has a much broader spectral line and shorter relaxation times, giving a lower SNR (Figure 8). Magnetization can be transferred between pools bidirectionally through direct interaction between spins, transfer of nuclei or direct chemical means. Under normal circumstances, magnetization transfer is the same in both directions.5Westbrook C. Roth C.K. Talbot J. MRI in Practice.4th edition. John Wiley & Sons, Inc., London2011Google ScholarFigure 8Model demonstrating the concepts underlying the phenomenon of magnetization transfer.View Large Image Figure ViewerDownload (PPT) Techniques employing MT saturate the magnetization in the bound pool, leaving the free pool mostly unaffected. This is possible due to the broad spectral line of the bound pool. It can be excited through the use of an “off-resonance” RF pulse (Figure 8). The saturation of the bound pool causes substantial attenuation of the magnetization. Consequently, there is little transfer of the magnetization to the free pool, with the effective longitudinal magnetization within it and its T1 relaxation time reduced as a consequence. Pulse sequences incorporating the use of “off-resonance” pulses can be designed to quantitate the effect of MT in different tissues.10Hajnal J.V. Baudouin C.J. Oatridge A. Young I.R. Bydder G.M. Desig and implementation of magnetization transfer pulse sequences for clinical use.J Comput Assist Tomogr. 1992; 16: 7-18Crossref PubMed Scopus (131) Google Scholar The free pool of protons (A) has a narrow spectral line, resonating at the Larmor frequency (ν0). RF pulses covering the frequencies, which are shown in pink (Figure 8), are able to excite the free pool. The “bound” pool (B) has a broad spectral line, while the subsequent application of RF irradiation at a frequency offset by Δν, shown in blue, can excite and saturate the pool without significantly affecting the free pool (pool A). Diffusion-weighted imaging (DWI) is an MR technique allowing quantification of water molecule movement. In the early 1990s, DWI was pioneered to detect acute cerebral ischemia.12Baird A.E. Warach S. Magnetic resonance imaging of acute stroke.J Cereb Blood Flow Metab. 1998; 18: 583-609Crossref PubMed Scopus (472) Google Scholar, 13Moseley M.E. Kucharczyk J. Mintorovitch J. et al.Diffusion-weighted MR imaging of acute stroke: correlation with T2 weighted and magnetic susceptibility-enhanced MR imaging in cats.AJNR Am J Neuroradiol. 1990; 11: 423-429PubMed Google Scholar Other indications include investigation for multiple sclerosis and brain tumors.14Larsson H.B. Thomsen C. Frederiksen J. Stubgaard M. Henriksen O. In vivo magnetic resonance diffusion measurement in the brain of patients with multiple sclerosis.Magn Reson Imaging. 1992; 10: 7-12Abstract Full Text PDF PubMed Scopus (158) Google Scholar, 15Kono K. Inoue Y. Nakayama K. et al.The role of diffusion-weighted imaging in patients with brain tumors.AJNR Am J Neuroradiol. 2001; 22: 1081-1088PubMed Google Scholar, 16Stadnik T.W. Chaskis C. Michotte A. et al.Diffuion-weighted MR imaging of intracerebral masses: comparison with conventional MR imaging and histologic findings.AJNR Am J Neuroradiol. 2001; 22: 969-976PubMed Google Scholar Water molecule diffusion follows the principles of Brownian motion. Thus, when unconstrained, water molecule movement is random and equal in all directions. This random movement is described as “isotropic”. However, motion of water molecules in structured environments is restricted due to their physical surroundings. In the brain, the microstructure within gray and white matter restricts water molecule movement. On average, water molecules tend to move parallel to white matter tracts, as opposed to perpendicular to them.17Chenevert T.L. Brunberg J.A. Pipe J.G. Anisotropic diffusion in human white matter: demonstration with MR techniques in vivo.Radiology. 1990; 177: 401-405Crossref PubMed Scopus (432) Google Scholar, 18Doran M. Hajnal J.V. Van Bruggen N. King M.D. Young I.R. Bydder G.M. Norma and abnormal white matter tracts shown by MR imaging using directional diffusion weighted sequences.J Comput Assist Tomogr. 1990; 14: 865-873Crossref PubMed Scopus (119) Google Scholar This motion is described as “anisotropic”, as it is not equal in all directions. The molecules’ motion in the x, y and z planes and the correlation between these directions is described by a mathematical construct, known as the diffusion tensor.19Basser P.J. Mattiello J. LeBihan D. MR diffusion tensor spectroscopy and imaging.Biophys J. 1994; 66: 259-267Abstract Full Text PDF PubMed Scopus (4430) Google Scholar, 20Basser P.J. Jones D.K. Diffusion-tensor MRI: theory, experimental design and data analysis—a technical review.NMR Biomed. 2002; 15: 456-467Crossref PubMed Scopus (1163) Google Scholar In mathematics, a tensor defines the properties of a three-dimensional ellipsoid. For the diffusion tensor to be determined, diffusion data in a minimum of six noncollinear directions are required. This process is known as diffusion tensor imaging (DTI). Figure 9 shows the graphical representation of a diffusion tensor, as a three-dimensional ellipsoid; the long axis represents the primary direction of motion.20Basser P.J. Jones D.K. Diffusion-tensor MRI: theory, experimental design and data analysis—a technical review.NMR Biomed. 2002; 15: 456-467Crossref PubMed Scopus (1163) Google Scholar Unconstrained, a water molecule would move randomly and equally in all directions, isotropic diffusion (A). The radius “r” of the spherical range of motion seen in Figure 9 defines the probability of motion in a given direction. Anisotropic diffusion will occur in an ordered environment, for example within the white matter, and will form an elliptical range of motion (B and C). Three eigenvalues, λ1, λ2, and λ3 and three eigenvectors v1, v2 and v3 define the shape and orientation of the ellipsoid, respectively (Figure 9), describing the magnitude and directions of the three major planes of the diffusion ellipsoid.21Mori S. van Zijl P. 2005, MR tractography using diffusion tensor imaging.in: Gillard J. Waldman A. Barker P.B. Clinical MR Neuroimaging. 1st edition. Cambridge University Press, 2005: 86-98Google Scholar During DTI, the tensor is calculated at each pixel location, allowing a map of diffusion to be produced, showing the magnitude and dominant direction of the process. When followed across a number of pixels, the dominant directions plot lines along which diffusion is most likely to occur. Practicing this technique is known as tractography, due to the theory that the likely diffusion of these paths represents the white matter tracts. DTI collects detailed information allowing insight into the microstructure found within an imaging voxel. Factors calculated include the mean diffusivity, degree of anisotropy, and direction of the diffusivities.22Jones D.K. Fundamentals of diffusion MR imaging.in: Gillard J. Waldman A. Barker P.B. Clinical MR Neuroimaging. 1st edition. Cambridge University Press, 2005: 54-85Google Scholar Mean diffusivity is a measure of displacement of water and also the presence of obstacles to movement at a cellular and subcellular level. Using differently weighted DWI images, a measure of diffusion can be calculated. The different images can be mapped to create an apparent diffusion coefficient (ADC) Image.23Mori S. Barker P.B. Diffusion magnetic resonance imaging: its principle and applications.Anat Rec. 1999; 257: 102-109Crossref PubMed Scopus (230) Google Scholar The ADC measures tissue water diffusivity dependent on the interactions between water molecules and their surrounding structural and chemical environment.24Le Bihan D. Breton E. Lallemand D. Grenier P. Cabanis E. Laval-Jeantet M. MR imaging of intravoxel incoherent motions: application to diffusion and perfusion in neurologic disorders.Radiology. 1986; 161: 401-407Crossref PubMed Scopus (3072) Google Scholar Fractional anisotropy (FA) and relative anisotropy (RA) are terms frequently used to describe the degree of anisotropy. Anisotropy relates to physical barriers, affected by characteristics, such as the density, orientation, size, and shape of nerve fibers within white matter tracts. However, myelination has been demonstrated not to be an essential component for anisotropy, though it certainly does contribute to the development of it, with nonmyelinated nerves also having the potential to exhibit anisotropy.25Beaulieu C. The basis of anisotropic water diffusion in the nervous system—a technical review.NMR Biomed. 2002; 15: 435-455Crossref PubMed Scopus (3479) Google Scholar The direction of the anisotropy, and therefore the fibers, can be plotted on color-coded two-dimensional maps (Figure 10), or otherwise by three-dimensional tractography. Various algorithms can be used to calculate the orientation of the major axonal fiber bundles using eigenvectors and eigenvalues.21Mori S. van Zijl P. 2005, MR tractography using diffusion tensor imaging.in: Gillard J. Waldman A. Barker P.B. Clinical MR Neuroimaging. 1st edition. Cambridge University Press, 2005: 86-98Google Scholar Three-dimensional expression of DTI data is one of the latest developments using this technique and may provide a better understanding into failings in brain connectivity. Clinical imaging at 3 T field strength instead of 1.5 T has advantages in diffusion-weighted imaging. The advantages are improved signal-to-noise ratio by 30–50%, improved contrast-to-noise ratio by up to 96% and reduced variability in ADC and FA by 34–52%.26Alexander A.L. Lee J.E. Wu Y.C. Field A.S. Comparison of diffusion tensor imaging measurements at 3.0 T versus 1.5 T with and without parallel imaging.Neuroimaging Clin N Am. 2006; 16: 299-309Abstract Full Text Full Text PDF PubMed Scopus (80) Google Scholar, 27Kuhl C.K. Textor J. 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Kooijman H. et al.Monitoring of gustatory stimulation of salivary glands by diffusion-weighted MR imaging: comparison of 1.5 T and 3 T.AJNR Am J Neuroradiol. 2007; 28: 1547-1551Crossref PubMed Scopus (31) Google Scholar, 30Lee C.E. Danielian L.E. Thomasson D. Baker E.H. Normal regional fractional anisotropy and apparent diffusion coefficient of the brain measured on a 3 T MR scanner.Neuroradiology. 2009; 51: 3-9Crossref PubMed Scopus (46) Google Scholar Cellular structures are not orientated in perfect symmetrical alignment homogeneously throughout the body, and thus the measurement of the diffusion of water molecules will be directionally dependent. This means that diffusion needs to be measured in several directions to obtain a rotationally invariant estimate of isotropic diffusion. Various experiments and modeling strategies have been employed to determine the minimal number of diffusion directions required to obtain an isotropic voxel, from which robust ADC and FA data can be derived, thus allowing a reasonable scan time for the patient with acquisition of reliable data. The minimal number of directions is reported as 20–30,31Correia M.M. Carpenter T.A. Williams G.B. Looking for the optimal DTI acquisition scheme given a maximum scan time: are more b-values a waste of time?.Magn Reson Imaging. 2009; 27: 163-175Abstract Full Text Full Text PDF PubMed Scopus (43) Google Scholar, 32Jones D.K. The effect of gradient samp