Basic MRI Physics Review Questions

MRI Pulse Sequences
MRI - Spatial Localization
MRI - Image Formation and Artifacts
MRI - Tissue Contrast
MRI - Chemical Shift
MRI - Diffusion-Weighted Imaging

Take these questions as a:

MRI Pulse Sequences

Back to section.

In the presence of a uniform magnetic field, hydrogen protons
1. Line up along the field and rotate around its axis
2. Line up along the field and precess around its axis
3. Remain oriented mostly randomly and precess around the field axis
4. Are not affected by the magnetic field
Protons will only have a slight predilection for orientation along the main magnetic field, being continually disturbed by molecular motion. They will precess (wobble) around the magnetic field axis, though. The net sum of the proton magnetization will, however, point along the main magnetic field and will not appear to precess (since the random precession of all the protons will cancel out).
MRI measures hydrogen magnetization
1. Along the main magnetic field
2. Perpendicular to the main magnetic field
The MRI scanner can only measure magnetization perpendicular to the main magnetic field (B0), as the magnetization precesses about B0.
In order to perform an MRI experiment, we _first_ need to
1. Apply a radiofrequency pulse in a different direction from B0
2. Apply a radiofrequency pulse along B0
3. Change the strength of B0
4. Apply a magnetic field gradient
We need to apply a radiofrequency pulse to rotate the proton magnetization so that it moves into the xy-plane - then we can measure the signal.
After a radiofrequency pulse, protons will
1. Dephase according to T2* and then recover along B0 according to T1
2. Dephase according to T1 and then recover along B0 according to T2*
3. Recover along B0 according to T1 and then dephase according to T2*
4. Recover along B0 according to T2* and then dephase according to T1
Proton dephasing in the xy plane occurs first, according to T2* characteristics, and longitudinal recovery along B0 is much slower, according to T1 characteristics.
The difference between T2 and T2* is
1. T2* represents the combination of static magnetic field inhomogeneity, which is T2, and varying molecular magnetic field inhomogeneity
2. T2 represents the combination of static magnetic field inhomogeneity, which is T2*, and varying molecular magnetic field inhomogeneity
3. T2* represents the combination of static magnetic field inhomogeneity and varying molecular magnetic field inhomogeneity, which is T2
4. T2 represents the combination of static magnetic field inhomogeneity and varying molecular magnetic field inhomogeneity, which is T2*
T2 is a characteristic of a particular tissue and represents dephasing related to varying molecular field inhomogeneities, related to the speed of molecular motion and vibration (also known as spin-spin relaxation).
A spin echo is formed by
1. Reversing the direction of B0
2. Flipping proton direction by a radiofrequency pulse
3. Applying a negative magnetic field gradient
A 180-degree radiofrequency pulse is applied to flip the direction of the protons; they continue to precess in the same direction along B0 but now rephase to the same point. A negative magnetic field gradient is what forms the gradient echo.
TR and TE in spin echo sequences refer to
1. TR: the spacing between successive 90-degree pulses; TE: time before the 90-degree pulse and the spin echo
2. TR: the spacing between successive 90-degree pulses; TE: the spacing between the 90-degree pulse and the 180-degree pulse
3. TR: time before the 90-degree pulse and the spin echo; TE: the spacing between successive 90-degree pulses
4. TR: the spacing between the 90-degree pulse and the 180-degree pulse; TE: the spacing between successive 90-degree pulses
The fast spin echo sequence alters the spin echo sequence by
1. Adding successive 90 degree pulses
2. Adding successive frequency-encoding gradients
3. Adding successive 180 degree pulses
4. Adding successive spoiler gradients
The fast spin echo uses multiple 180 degree pulses after a single 90-degree pulse to obtain multiple spin echoes from each excitation. Successive frequency-encoding gradients are used in echo planar imaging.
Compared with spin echo, fast spin echo sequences have
1. Increased T2 weighting but decreased susceptibility
2. Increased T2 weighting and increased susceptibility
3. Decreased T2 weighting but increased susceptibility
4. Decreased T2 weighting and decreased susceptibility
FSE sequences are more T2 weighted because many echoes are acquired at late echo times; susceptibility is decreased because of decreased time for dephasing between the successive 180-degree pulses.
Compared with spin echo (SE), the gradient echo (GRE) sequence
1. Has a 180-degree pulse and demonstrates T2* dephasing
2. Has no 180-degree pulse and demonstrates T2* dephasing
3. Has a 180-degree pulse and demonstrates T2 dephasing
4. Has no 180-degree pulse and demonstrates T2 dephasing
Gradient echo sequences use the free induction decay signal from the initial excitation pulse; they do not form a spin echo using a 180-degree pulse. Thus, the dephasing occurs according to T2* rather than T2.
Compared with SE, GRE sequences use
1. Longer TR and longer TE
2. Shorter TR and longer TE
3. Longer TR and shorter TE
4. Shorter TR and shorter TE
Because of T2* dephasing, GRE sequences must use much shorter TEs. Because of this, GRE sequences can use much shorter TRs.
In GRE sequences, the trade-off with _decreasing_ flip angle for the excitation pulse is
1. Increased T1 weighting but decreased signal
2. Decreased T1 weighting but improved signal
3. Increased T2* weighting but decreased signal
4. Decreased T2* weighting but improved signal
Smaller flip angles decrease the T1 weighting of the sequence and are used to preserve longitudinal magnetization (signal) for future repetitions of the sequence.
Traditional GRE sequences
1. Use long TR to reduce effects from residual transverse magnetization
2. Use spoiler gradients to reduce effects from residual transverse magnetization
3. Preserve transverse magnetization for successive repetitions
Compared to traditional GRE sequences, balanced steady-state free precession (SSFP) sequences
1. Use 180-degree pulses to form spin echoes
2. Retain transverse magnetization to form spin echoes
3. Use spoiler gradients to reduce transverse magnetization
4. Use alternating magnetic field gradients to traverse k-space
Balanced SSFP sequences keep transverse magnetization, which forms spin echoes after successive excitation pulses; these contribute to signal along with the FID signal. 180-degree pulses are used in spin echo sequences; spoiler gradients are used in traditional GRE; and alternating gradients are used in echo planar imaging.
Balanced SSFP sequences are used for cardiac and vascular imaging because they are
1. Flow sensitive
2. Flow insensitive
3. Null blood signal
4. Employ gadolinium contrast
Echo planar imaging (EPI) is most commonly used in
1. Vascular imaging
2. Routine T2-weighted imaging
3. Post-contrast T1-weighted imaging
4. Diffusion-weighted imaging
DWI requires very fast sequences to freeze patient motion, and EPI is extremely fast.
Echo planar imaging forms echoes using
1. Successive 90-degree pulses
2. Successive alternating 90-degree pulses
3. Successive 180-degree pulses
4. Successive alternating gradients
EPI initially forms a spin echo and then uses alternating gradients to create multiple gradient echoes within that spin echo.

MRI - Spatial Localization

Back to section.

Spatial localization in MRI primarily relies on
1. Distance to the receiving coil
2. Distance from the transmission coil
3. Varying magnetic field across the patient
4. Tomographic reconstruction
Gradient coils create varying magnetic fields across the patient, causing changes in precession frequency.
Slice-selection is performed by
1. Turning on a gradient coil prior to excitation
2. Turning on a gradient coil during excitation
3. Turning on a gradient coil prior to readout
4. Turning on a gradient coil during readout
Slice-selection is accomplished by turning on a gradient coil during excitation (and refocusing, if using spin echo). Tuning the excitation pulse to the frequency of the relevant slice then only excites protons in that slice.
In order to select a slice for excitation, the MR scanner
1. Tunes the frequency of the excitation pulse
2. Tunes the phase of the excitation pulse
3. Tunes the magnetic field of the excitation pulse
4. Tunes the spacing between the excitation and refocusing (180-degree) pulses
The slice-select gradient sets up a magnetic field gradient across the patient, changing the Larmor frequency of the protons in the entire body. By matching the excitation pulse to the frequency of protons at a particular position, the scanner can excite a thin slice of the body.
The Fourier transform changes signals in what manner
1. Samples over time to amplitudes separated by phase
2. Samples over time to frequencies separated by amplitude
3. Samples over time to amplitudes separated by frequency
4. Set of frequencies to samples over time
5. Continuous signal to set of samples over time
The Fourier transform separates a signal in time into its component frequencies, reporting the amplitudes (amount) for each frequency.
When digitizing a signal, the sampling (measurement) frequency should be set based on
1. The maximum amplitude in the signal
2. The maximum phase shift in the signal
3. The maximum frequency in the signal
4. The desired frequency resolution
5. It doesn't matter
By the Nyquist sampling theorem, the sampling frequency must be at least twice the maximum frequency in the signal to prevent aliasing. Frequency resolution is determined by the number of samples.
If a signal is undersampled, aliasing will result and cause
1. Amplitude misregistration
2. Frequency misregistration
3. Phase misregistration
4. Poor resolution
5. Noise
Aliasing is wrap of too-high frequencies onto other frequencies in the resulting Fourier transform (or image, in the case of MRI). As noted, resolution is determined by matrix size.
Within-slice signal localization in MRI is performed with
1. Tomography
2. Iterative reconstruction
3. Fourier transform
4. Selective excitation
The Fourier transform is applied to a matrix created with frequency- and phase-encoding to generate the image. MR does not presently use tomography (although the initial paper did describe a tomographic technique using only frequency encoding!) Selective excitation is used for slice selection.
The strength of the frequency-encoding gradient
1. Is varied multiple times to generate K-space
2. Is irrelevant for image appearance
3. Alters the bandwidth of the signal
4. Alters the tissue contrast
Varying the frequency-encoding gradient changes the signal bandwidth (range of frequencies present in the image) since it changes the range of magnetic field the protons are exposed to. This has implications in signal-to-noise and chemical shift effects. Multiple applications of different strength phase-encoding gradients are used to fill K-space. The gradient strength does not change the T1/T2 weighting.
The phase-encoding gradient is turned on
1. During excitation
2. During slice-selection
3. During frequency-encoding
4. Prior to readout
5. During readout
The phase-encoding gradient is turned on prior to readout - and then turned off. This imparts a fixed phase shift to the protons. Frequency-encoding is turned on during readout. The slice-select gradient is turned on during excitation but not readout.
The Fourier transform can directly measure phase shifts in an MR signal
1. True
2. False
The Fourier transform can measure only a single phase shift at each frequency. That is why multiple phase-encoding steps are necessary (one for each pixel of resolution); the Fourier transform can detect a changing phase shift.
In MR imaging, matrix size determines
1. Field of view
2. Aliasing
3. Resolution
4. Bandwidth
The matrix size is directly related to phase- and frequency-resolution (i.e. the number of different phases or frequencies which can be detected), which directly translates into image resolution. Field of view is determined by sampling frequency - the spacing between phase or frequency samples - which determines the maximum frequency measurable (and thus the maximum gradient shift measurable). Thus, spacing affects aliasing as well if sampling < Nyquist limit. Bandwidth is determined by gradient strength.
The most signal in K-space is present
1. In the center
2. In the periphery
3. Along the frequency-encoding axis
4. Along the phase-encoding axis
The most signal in K-space is present at the center. The periphery loses signal because of dephasing.
The center of K-space contains
1. Frequency information
2. Degree of T2 weighting
3. Tissue contrast
4. Temporal resolution
5. Spatial resolution
The signal in the center of K-space determines the contrast in the image, i.e. the brightness of large objects. The periphery determines the resolution, i.e. the tiny details. Temporal resolution is generally determined by the number of phase-encoding steps (which take the longest time).

MRI - Image Formation and Artifacts

Back to section.

SNR in MRI is improved by increasing:
1. Resolution
2. Bandwidth
3. Gradient strength
4. Acquisition time
Intuitively, the more time spent actually acquiring the signal, the more signal we obtain.
Increasing bandwidth causes:
1. Worsened chemical shift artifact
2. Worsened SNR
3. Worsened metal hardware artifacts
4. Longer echo acquisition time
Increased BW improves chemical shift and metal artifacts. The echo is faster to acquire (because we are acquiring higher frequencies). However, we suffer from worse SNR because we are letting in more frequencies (and noise is distributed across all frequencies).
Which of the following is true about reducing phase-encoding steps:
1. Increasing k-space spacing -> worse resolution
2. Dropping peripheral k-space lines (dropping scan percentage) -> smaller FOV
3. Half-Fourier acquisition -> worse resolution
4. Worse SNR
Again, the less time we spend acquiring the signal, the less signal we get. Increasing spacing in k-space decreases FOV. Dropping peripheral lines in k-space yields worse resolution. Half-Fourier acquisition gives worse SNR but same resolution (k-space is symmetric so we can just mathematically reconstruct the rest of the data).
The spacing of lines in k-space corresponds to:
1. Resolution
2. Number of excitations
3. Matrix size
4. Field of view
Spacing of lines in k-space is inversely related to FOV. Resolution is related to the number of lines in k-space we acquire (how far we go to the periphery of k-space). Matrix size is also related to the number of lines in k-space but not FOV. Excitations are not related at all.
Aliasing refers to the mathematical phenomenon of:
1. Frequency or phase misregistration because of signal undersampling
2. Pixel misregistration related to gyromagnetic differences
3. Pixel blurring caused by lack of sampling high frequencies
4. Pixel misregistration because of patient motion
Aliasing is the phenomenon of misinterpreting the phase or frequency of a signal because of inadequate sampling resolution. This is most often seen with rectangular FOV - we do not sample k-space closely enough and so we get a smaller FOV. The Fourier transform does not know what to do with the voxels outside of the narrow FOV, so it wraps them in. Aliasing is also evident in phase-contrast MRI or spectral Doppler ultrasound.
Radiofrequency contamination artifact causes:
1. "Ghosting" of multiple copies of structures
2. Diagonal lines across the image
3. Random noise across the image
4. Line(s) parallel to the frequency-encoding axis
5. Line(s) parallel to the phase-encoding axis
RF contamination creates lines parallel to the phase-encoding axis - i.e. at a constant frequency. Ghosting refers to motion artifacts (along the phase-encoding axis). Diagonal lines refers to the 'zebra' or 'spike' artifact. Random noise is always present in an image, related to SNR.
Radiofrequency contamination artifact is best addressed by:
1. Checking or replacing the receiver coils
2. Checking or replacing the transmit coils
3. Shimming the magnet
4. Checking or replacing room and door shielding
RF contamination is often related to someone inadvertently leaving the door open! But it can also be caused by damage to the RF shielding (copper mesh, Faraday cage) around the room or inappropriate hardware left in the room.
Motion artifact is caused by:
1. Long time in between acquisition of points in k-space along the frequency-encoding direction
2. Long time in between acquisition of points in k-space along the phase-encoding direction
3. Frequency shifts in protons related to velocity
4. Dephasing of protons related to velocity
Motion artifact is caused by changes in the position of the patient - or vascular structures - in between phase-encoding steps.
You are attempting to perform a liver MRI to evaluate for metastatic disease, and a copy ('ghost') of the aorta appears over segment 3. What is the best way to remove this artifact?
1. Swap phase and frequency encoding directions
2. Phase oversampling
3. Increase bandwidth
4. Give intravenous contrast
5. Use an inversion-recovery sequence
This pulsation artifact is caused by motion of the aortic wall and blood flow within the aorta. As a motion artifact, it occurs in the phase-encoded direction. By switching the axes, it will now occur in the left-right axis and not overlap the liver.

MRI - Tissue Contrast

Back to section.

MRI employs differences in what physical phenomena to obtain tissue contrast:
1. Gyromagnetic ratio
2. Spin-spin and spin-lattice relaxation constants
3. Carbon and nitrogen abundance
4. Magnetic field inhomogeneity
Tissues differ in their mobile proton density, spin-spin (T2), and spin-lattice (T1) relaxation constants.
In order to measure recovery in the longitudinal axis (T1 relaxation), an MR scanner measures:
1. Signal in the longitudinal axis (z-axis, along B0)
2. Signal in the transverse plane (xy-plane)
3. Proton precession frequency
4. Local magnetic field
The MR receiver coil can only measure signal in the transverse plane. This is why the signal needs to be flipped from the z-axis into the xy-plane by a 90-degree pulse.
The MR imaging parameter that determines how much T1 (longitudinal) recovery is allowed to occur is the
1. TR
2. TE
3. Bandwidth (BW)
4. Number of excitations
5. Echo train length (ETL)
The TR, or repetition time, is the time between 90 degree pulses. This represents the amount of time signal in the longitudinal axis is allowed to recover.
A tumor that enhances with gadolinium contrast is bright on T1-weighted images because
1. Gadolinium chelates slow T1 recovery of water molecules by interacting with the water and changing the rotational speeds of the hydrogen atoms
2. Gadolinium has a short T1 value
3. The hydrogen atoms in the chelation groups of the contrast agent have a short T1 value
4. Gadolinium chelates speed T1 recovery of water molecules by interacting with the water and changing the rotational speeds of the hydrogen atoms
The gadolinum and chelation groups are not imaged with MR. Gadolinium chelate complexes interact with water and slow the rotation speeds of the hydrogen atoms, thus bringing them closer to the resonant frequency of hydrogen, allowing for faster energy exchange and thus quicker T1 relaxation.
Increasing B₀ affects the T1 value of water in what way?
1. Longer T1
2. Shorter T1
3. Does not change T1
4. Only changes T1 of water-protein mixtures
Increasing B0 makes T1 longer because fewer water protons are moving at the correct resonant frequency, thus slowing energy exchange.
T2-weighted images are often referred to as "water-sensitive" because
1. The presence of water alters the rotation speeds of the normal tissue
2. Water molecules have very long (slow) T2 relaxation
3. Water molecules have very short (long) T2 relaxation
4. Trick question - T2-weighted images are not water sensitive
5. T2-weighted images are only sensitive to water-gadolinium complexes
Water molecules have very long (slow) T2 relaxation because of high molecular rotation speeds. This leads to bright signal on T2-weighted images, which are thus often referred to as "water-sensitive."
It is difficult to perform MR imaging of mummies because
1. Solid (dry) tissues have very long T1
2. Solid (dry) tissues have very long T2
3. Solid (dry) tissues have very short T1
4. Solid (dry) tissues have very short T2
Solid matter has very short T2 because of slow to no molecular rotation, which causes magnetic field inhomogeneities within the tissue and speeds proton dephasing. Mummies can be imaged with T1-weighted sequences, however.
Inversion-recovery (IR) sequences are helpful to
1. Improve T2 weighting
2. Improve signal-to-noise (SNR)
3. Shorten imaging time
4. Improve tissue contrast
Tissue contrast is improved by (a) nulling of signal from a particular type of tissue (e.g. water in FLAIR or fat in STIR) and by (b) improving T1 weighting. T2 weighting is unaffected and still determined by TE. SNR is worse because of the greater T1 weighting and thus poorer signal recovery (especially with STIR). Imaging time is longer because we have to wait for the extra TI.
Reasonable parameters for a T1-weighted image would be
1. TR 500 ms, TE 20 ms
2. TR 2800 ms, TE 90 ms
3. TR 6000 ms, TE 160 ms, TI 2000 ms
4. TR 2800 ms, TE 20 ms
T1-weighting is driven by short TR, giving less time for T1 recovery and thus more differences in signal based on T1 characteristics. You want a short TE to minimize T2 effects. Long TR minimizes T1 effects; with a long TE, that gives T2 weighting, while with a short TE that gives proton density weighting.
To null the signal from a particular tissue with an IR sequence, TI should be chosen based on
1. The proton density of the tissue
2. The T1 value of the tissue
3. The T2 value of the tissue
4. The T2* value of the tissue
5. The precession frequency of the tissue
Inversion recovery is based on the T1 characteristics of the tissue. TI is typically near the T1 value of the tissue, although the exact best TI depends on the TR of the sequence as well. A saturation pulse tuned to the precession frequency of a tissue is used in chemical fat suppression.

MRI - Chemical Shift

Back to section.

Protons in different molecules differ in all of the following ways except
1. T1
2. T2
3. Gyromagnetic ratio
4. Precession frequency
The gyromagnetic ratio is a physical constant for the proton. However, the precessional frequency of the proton will differ based on the molecular electric field (often referred to as magnetic shielding). T1 and T2 characteristics differ as well, related to molecular rotation speeds.
Chemical shift artifact causes misregistration of fat and water signal in which of the following ways:
1. Fat signal is collected in a separate echo from water signal
2. Fat signal is shifted along the frequency-encoding axis
3. Fat signal is shifted along the phase-encoding axis
4. Fat signal is shifted along the slice-selection axis
Because chemical shift alters the frequency of proton precession, the signal appears at a different location along the frequency-encoding axis.
Increasing bandwidth has what effect on chemical shift artifacts:
1. Worsens chemical shift artifacts
2. Improves signal misregistration but not destructive interference (phase cancellation)
3. Improves signal misregistration and destructive interference (phase cancellation)
4. Worsens signal misregistration but does not affect destructive interference
5. No effect on chemical shift artifact
Bandwidth affects the size of frequencies assigned to the same pixel. Thus, it improves signal misregistration because of chemical shift effects. However, it has no effect on the phase differences between the signals themselves; thus, they will still cancel out when they are out of phase.
Fat-water phase differences in an MR image are determined by what imaging parameter
1. TR in a spin echo (SE) sequence
2. TE in a spin echo (SE) sequence
3. TR in a gradient echo (GRE) sequence
4. TE in a gradient echo (GRE) sequence
5. TR in either SE or GRE sequence
6. TE in either SE or GRE sequence
The amount of phase difference depends on the time used to measure the actual echo (i.e. TE). These do not occur in spin echo sequences because the 180-degree pulse rewinds the phase shifts.
The Dixon method of fat suppression relies on
1. Obtaining a water-only image by varying TE
2. Obtaining a fat-suppressed image with an inversion pulse
3. Mathematically calculating a water-only image by acquiring two echoes
4. Employing a saturation pulse based on the precessional frequency of fat
The Dixon method acquires in-phase and out-of-phase images and then adds and subtracts them to obtain fat-only and water-only images. The water-only image cannot be directly measured. Fat suppression with an IR pulse is referred to as STIR. Fat suppression with a pulse based on the precessional frequency of fat is referred to as chemical fat suppression or fat saturation.

MRI - Diffusion-Weighted Imaging

Back to section.

Diffusion-weighted imaging measures the:
1. Motion of water molecules
2. Motion of cells
3. Chemical composition of tissue
4. Ratio of water to fat
Signal on DWI is related to diffusion (motion) of water molecules in the voxel. The more the water molecules move (faster diffusion), the lower the signal.
Diffusion is typically NOT restricted by
1. Intracellular water
2. Extracellular water
3. Pus
4. Tumor cells
Water tends to diffuse freely in the extracellular compartment, where it is not hindered by cell membranes.
The diffusion weighting in DWI images is created by means of
1. Two balanced gradients spaced in time
2. Triphasic flow compensation gradients
3. An inversion pulse
4. Two inversion pulses
Two gradients are applied, with the second one the inverse of the first. These should cause no net phase shift in stationary tissue; any protons that have moved will accumulate a phase shift.
The most common pulse sequence used for DWI is
1. Standard spin echo (SE)
2. Fast spin echo (FSE)
3. Half-Fourier fast spin echo
4. Gradient echo (GRE)
5. Steady-state free precession (SSFP)
6. Echo-planar imaging (EPI)
To eliminate patient motion (which will overwhelm the diffusion phenomenon), we must use the fastest imaging sequence possible - echo-planar. Echo planar images are subject to severe field inhomogeneity artifacts, limiting usefulness near air or bone interfaces.
Typical pathology as seen on DWI and ADC images is
1. Bright on DWI and bright on ADC
2. Bright on DWI and dark on ADC
3. Dark on DWI and bright on ADC
4. Dark on DWI and dark on ADC
Typical pathology (i.e. true restricted diffusion) will be bright on DWI images and dark on the ADC map, reflecting a low diffusion coefficient.
"False-positive" findings on DWI are often attributed to
1. T1 effects
2. T2 effects
3. Anisotropy
4. Poor signal to noise
T2 shine-through can cause bright signal on DWI images because of the inherent (long) T2 properties of the tissue, irrespective of the diffusivity. This is why it is important to verify findings with ADC maps.
ADC maps negate T2 shine-through by
1. Using multiple directions of diffusion gradients
2. Using multiple intensities of diffusion gradients
3. Averaging multiple acquisitions of diffusion
4. Using a different pulse sequence to acquire diffusion
By using multiple different b-values, we can calculate the slope of the signal loss with increasing diffusion weighting. This yields the apparent diffusion coefficient (ADC), which is independent of T2 effects.
Utilizing a single direction of diffusion, white matter tracts would appear to
1. Uniformly restrict diffusion because of their organized structure
2. Uniformly show fast diffusion because of their organized structure
3. Heterogeneously restrict diffusion depending on their orientation
4. Uniformly show dark signal because of their fat content
White matter tracts are highly organized and allow water to diffuse primarily along the direction of the axons. Measuring only one direction of diffusion will show fast diffusion when the tracts are oriented in that direction and restricted diffusion when they are oriented in other directions. This is why measuring several different directions of diffusion is important.