Table of Contents
Cover
Title Page
Copyright
Preface
ABR study guide topics
Chapter 1: Production of net magnetization
1.1 Magnetic fields
1.2 Nuclear spin
1.3 Nuclear magnetic moments
1.4 Larmor precession
1.5 Net magnetization
1.6 Susceptibility and magnetic materials
Chapter 2: Concepts of magnetic resonance
2.1 Radiofrequency excitation
2.2 Radiofrequency signal detection
2.3 Chemical shift
Chapter 3: Relaxation
3.1
T1
relaxation and saturation
3.2
T2
relaxation,
T2
* relaxation, and spin echoes
Chapter 4: Principles of magnetic resonance imaging – 1
4.1 Gradient fields
4.2 Slice selection
4.3 Readout or frequency encoding
4.4 Phase encoding
4.5 Sequence looping
Chapter 5: Principles of magnetic resonance imaging – 2
5.1 Frequency selective excitation
5.2 Composite pulses
5.3 Raw data and image data matrices
5.4 Signal-to-noise ratio and tradeoffs
5.5 Raw data and
k
-space
5.6 Reduced
k
-space techniques
5.7 Reordered
k
-space filling techniques
5.8 Other
k
-space filling techniques
5.9 Phased-array coils
5.10 Parallel acquisition methods
Chapter 6: Pulse sequences
6.1 Spin echo sequences
6.2 Gradient echo sequences
6.3 Echo planar imaging sequences
6.4 Magnetization-prepared sequences
Chapter 7: Measurement parameters and image contrast
7.1 Intrinsic parameters
7.2 Extrinsic parameters
7.3 Parameter tradeoffs
Chapter 8: Signal suppression techniques
8.1 Spatial presaturation
8.2 Magnetization transfer suppression
8.3 Frequency-selective saturation
8.4 Nonsaturation methods
Chapter 9: Artifacts
9.1 Motion artifacts
9.2 Sequence/Protocol-related artifacts
9.3 External artifacts
Chapter 10: Motion artifact reduction techniques
10.1 Acquisition parameter modification
10.2 Triggering/Gating
10.3 Flow compensation
10.4 Radial-based motion compensation
Chapter 11: Magnetic resonance angiography
11.1 Time-of-flight MRA
11.2 Phase contrast MRA
11.3 Maximum intensity projection
Chapter 12: Advanced imaging applications
12.1 Diffusion
12.2 Perfusion
12.3 Functional brain imaging
12.4 Ultra-high field imaging
12.5 Noble gas imaging
Chapter 13: Magnetic resonance spectroscopy
13.1 Additional concepts
13.2 Localization techniques
13.3 Spectral analysis and postprocessing
13.4 Ultra-high field spectroscopy
Chapter 14: Instrumentation
14.1 Computer systems
14.2 Magnet system
14.3 Gradient system
14.4 Radiofrequency system
14.5 Data acquisition system
14.6 Summary of system components
Chapter 15: Contrast agents
15.1 Intravenous agents
15.2 Oral agents
Chapter 16: Safety
16.1 Base magnetic field
16.2 Cryogens
16.3 Gradients
16.4 RF power deposition
16.5 Contrast media
Chapter 17: Clinical applications
17.1 General principles of clinical MR imaging
17.2 Examination design considerations
17.3 Protocol considerations for anatomical regions
17.4 Recommendations for specific sequences and clinical situations
References and suggested readings
Index
End User License Agreement
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Guide
Cover
Table of Contents
Preface
Begin Reading
List of Illustrations
Chapter 1: Production of net magnetization
Figure 1.1 A rotating nucleus (spin) with a positive charge produces a magnetic field known as the magnetic moment oriented parallel to the axis of rotation (a). This arrangement is analogous to a bar magnet in which the magnetic field is considered to be oriented from the south to the north pole (b).
Figure 1.2 Microscopic and macroscopic pictures of a collection of spins in the absence of an external magnetic field. In the absence of a magnetic field, the spins have their axes oriented randomly (microscopic picture, left side of figure). The vector sum of these spin vectors is zero (macroscopic picture, right side).
Figure 1.3 Inside a magnetic field, a proton precesses or revolves about the magnetic field. The precessional axis is parallel to the main magnetic field,
B
0
. The
z
component of the spin vector (projection of the spin onto the
z
axis) is the component of interest because it does not change in magnitude or direction as the proton precesses. The
x
and
y
components vary with time at a frequency ω
0
proportional to
B
0
as expressed by equation (1.1).
Figure 1.4 Zeeman diagram. In the absence of a magnetic field (left side of figure), a collection of spins will have the configurations of
z
components equal in energy so that there is no preferential alignment between the spin-up and spin-down orientations. In the presence of a magnetic field (right side), the spin-up orientation (parallel to
B
0
) is of lower energy and its configuration contains more spins than does the higher-energy spin-down configuration. The difference in energy Δ
E
between the two levels is proportional to
B
0
.
Figure 1.5 Microscopic (a) and macroscopic (b) pictures of a collection of spins in the presence of an external magnetic field. Each spin precesses about the magnetic field. If a rotating frame of reference is used with a rotation rate of ω
0
, the collection of protons appears stationary. Whereas the
z
components are one of two values (one positive and one negative), the
x
and
y
components can be any value, positive or negative. The spins will appear to track along two cones, one with a positive
z
component and one with a negative
z
component. Because there are more spins in the upper cone, there will be a nonzero vector sum
M
0
, the net magnetization. It will be of constant magnitude and parallel to
B
0
.
Chapter 2: Concepts of magnetic resonance
Figure 2.1 Energy absorption (microscopic). The difference in energy Δ
E
between the two configurations (spin up and spin down) is proportional to the magnetic field strength
B
0
and the corresponding precessional frequency ω
0
, as expressed in Equation (2.1). When energy at this frequency is applied, a spin from the lower-energy state is excited to the upper-energy state. Also, a spin from the upper-energy state is stimulated to give up its energy and relax to the lower-energy state. Because there are more spins in the lower-energy state, there is a net absorption of energy by the spins in the sample.
Figure 2.2 Energy absorption (macroscopic). In a rotating frame of reference, the RF pulse broadcast at the resonant frequency ω
0
can be treated as an additional magnetic field
B
1
oriented perpendicular to
B
0
. When energy is applied at the appropriate frequency, the spins absorb it and
M
rotates into the transverse plane. The initial direction of rotation is perpendicular to both
B
0
and
B
1
. The amount of resulting rotation of
M
is known as the pulse flip angle.
Figure 2.3 (a) Free induction decay, real and imaginary. The response of the net magnetization M to an RF pulse as a function of time is known as the free induction decay (FID). Its amplitude is proportional to the amount of transverse magnetization generated by the pulse. The FID is maximized when using a 90° excitation pulse. (b) Fourier transformation of (a), magnitude and phase. The Fourier transformation is used to convert the digital version of the MR signal (FID) from a function of time to a function of frequency. Signals measured with a quadrature detector are displayed with the transmitter (reference) frequency
ω
TR
in the middle of the display. The Nyquist frequencies
ω
NQ
below and above
ω
TR
are the minimum and maximum frequencies of the frequency display, respectively. For historical reasons, frequencies are plotted with lower frequencies on the right side and higher frequencies on the left side of the display.
Figure 2.4 Planar (a) and circular (b) representations of a time-varying wave. The amplitude (A) is the maximum deviation of the wave from its mean value. The period (B) is the time required for completion of one complete cycle of the wave. The frequency of the wave is the reciprocal of the period. The phase or phase angle of the wave (C) describes the shift in the wave relative to a reference (a second wave for the planar representation, horizontal axis for circular representation). The two plane waves displayed in (a) have the same amplitude and period (frequency) but have a phase difference of π/4, or 90°.
Figure 2.5 Spectrum of water and fat at 1.5 T (a) and 3.0 T (b). The resonant frequencies for water and fat are separated by approximately 3.5 ppm, which corresponds to an absolute frequency difference of 220 Hz for a 1.5 T magnetic field (63 MHz) or 450 Hz at a magnetic field of 3.0 T (126 MHz).
Figure 2.6 Precession of fat and water protons. Because of the 3.5 ppm frequency difference, a fat proton precesses at a slower frequency than does a water proton. In a rotating frame at the water resonant frequency, the fat proton cycles in and out of phase with the water proton. Following the excitation pulse, the two protons are in phase (a). After a short time, they will be 90° out of phase (b), then 180° out of phase (c, also called “opposed phase”). Then −90° out of phase (d) and back in phase (e). The contribution of fat to the total signal fluctuates and depends on when the signal is detected. At 1.5 T, the in-phase times are 0 ms (a), 4.5 ms (e), 9 ms and so on (not shown), while the opposed-phase times are 2.25 ms (c), 6.7 ms and so on (not shown). At 3.0 T, the times are one half of the times at 1.5 T.
Chapter 3: Relaxation
Figure 3.1
T1
relaxation curve. Immediately following a 90° RF pulse, there is no longitudinal magnetization. A short time later, longitudinal magnetization will be observed as the protons release their energy through
T1
relaxation. Gradually, as more protons release their energy, a larger fraction of
is reestablished. Eventually,
will be restored completely. The change of
with time
follows an exponential growth process as described by equation (3.1). The time constant for this process is
T1
, the spin-lattice relaxation time, and is the time when
has returned to 63% of its original value.
Figure 3.2 Following a 90° RF pulse, longitudinal magnetization is regenerated through
T1
relaxation. If the time between successive RF pulses
is insufficient for complete recovery of
M
, only
will be present at the time of the next RF pulse (a). If time
elapses again, only
will be present (b).
will be smaller than
, but the difference will be less than the difference between
M
and
.
Figure 3.3 Saturation. If RF pulses are applied faster than the
T1
relaxation processes can dissipate the energy, the spin populations equalize between the two energy levels. As a result, there is no difference in the number of spins and no net magnetization.
Figure 3.4 Spin– spin relaxation. (a) Two water molecules, with one hydrogen atom on one molecule having absorbed RF energy and being excited (spin down). (b) If the molecules are in close proximity, the energy can be transferred from the first water molecule to a hydrogen atom on the second water molecule.
Figure 3.5 (a) A rotating frame slower than
is assumed for this Figure Net magnetization
M
(arrow) is oriented parallel to
(not shown) prior to an RF pulse (1). Following a 90° RF pulse, the spins precess initially in phase in the transverse plane (2). Due to inter-and intramolecular interactions, the spins begin to precess at different frequencies (dashed arrow, faster; dotted arrow, slower) and become asynchronous with each other (3). As more time elapses (4,5), the transverse coherence becomes smaller until there is complete randomness of the transverse components and no coherence (6). (b) Plot of relative
component as a function of time. The numbers correspond to the expected
component from (a). The change in
with time follows an exponential decay process as described by equation (3.3). The time constant for this process is the spin-spin relaxation time
and is the time when
has decayed to 37% of its original value.
Figure 3.6 A rotation frame slower than
is assumed for this Figure Net magnetization
M
(arrow) is oriented parallel to
(not shown) prior to an RF pulse (1). Application of a 90° RF pulse rotates
M
into the transverse plane (2). Due to the
relaxation processes, the protons become asynchronous with each other during time
t
(3). Application of a 180° RF pulse causes the protons to reverse their phase relative to the transmitter phase. The protons that precessed most rapidly are farthest behind (dashed arrow), while the slowest protons are in front (dotted arrow) (4). Allowing time t to elapse again allows the protons to regain their phase coherence in the transverse plane (5), generating a signal in the receiver coil known as a spin echo. The loss in magnitude of the reformed coherence relative to the original coherence (2) is due to irreversible processes (i.e., true spin–spin or
T2
relaxation). Equation (3.3) describes the decay of
if
T2
is used instead of
.
Chapter 4: Principles of magnetic resonance imaging – 1
Figure 4.1 A gradient is a linear distortion of the primary magnetic field, centered at the magnet isocenter. A 40 mT m
−1
gradient superimposed on a 1.5 T magnet field will produce a total magnetic field variation of 1480 to 1520 mT. For a 500 mm distance, the variation will be 1490 to 1510 mT.
Figure 4.2 For a given range (bandwidth) of frequencies included in the RF pulse, the slice thickness desired is determined by the slice-selection gradient amplitude. The user interface typically allows variation of the slice thickness, which is achieved by increasing or decreasing the slice-selection gradient amplitude, as appropriate.
Figure 4.3 Slice-selection process. In the presence of a gradient
, the total magnetic field that a proton experiences and its resulting resonant frequency depend on its position according to equation 4.2. Tissue located at position
will absorb RF energy broadcast with a center frequency
. Each position will have a unique resonant frequency. The slice thickness
is determined by the amplitude (magnitude) of
and by the bandwidth of transmitted frequencies
.
Figure 4.4 Images in standard slice orientations: sagittal, coronal, and transverse or axial. For transverse images, two view directions are possible: cranial and caudal. Image annotations are based on patient axes.
Figure 4.5 Readout process. Following excitation, each spin within the excited volume (slice) precesses at the same frequency. During detection of the echo, a gradient
is applied, causing a variation in the frequencies for the spins generating the echo signal. The frequency of precession
for each spin depends on its position
according to equation (4.2). Frequencies measured from the echo are mapped to the corresponding position.
Figure 4.6 In any image, one of the directions visualized is the readout direction and the other is the phase-encoding direction. A proton located at the edge of the FOV in the readout direction precesses at the Nyquist frequency
above or below the transmitter frequency
. Changing the FOV of the image changes the spatial resolution (millimeters per pixel) but not the frequency resolution (hertz per pixel).
Figure 4.7 For a given range (bandwidth) of frequencies that are measured in the signal, the desired
FOV
is determined by the readout gradient amplitude. The user interface typically allows variation of the
FOV
, which is achieved by increasing or decreasing the readout gradient amplitude, as appropriate.
Figure 4.8 Concept of phase encoding. Prior to application of
, all spins precess at the same frequency. When
is applied, a spin increases or decreases its precessional frequency, depending on its position
. A spin located at
experiences no effect from
and no change in frequency or phase
. A spin located at
precesses faster while
is applied. Once
is turned off, the spin precesses at its original frequency but is ahead of the reference frequency (dashed curve); that is, a phase shift
has been induced to the proton by
. A spin located at
decreases its frequency while
is applied. Once
is turned off, it precesses at its original frequency but is behind the reference by a phase shift of
.
Figure 4.9 Phase-encoding process. A spin at the edge of the
FOV
in the phase-encoding direction undergoes 90° of phase change
from one phase-encoding step to the next. Each point within the
FOV
undergoes progressively less phase change for the same gradient amplitude. A spin at isocenter never experiences any phase change. The change in gradient amplitude (
in this example) from one phase-encoding step to the next is inversely proportional to the
FOV
.
Figure 4.10 Two-dimensional slice loop structures. Three slice loop structures are commonly used. (a) Traditional multislice looping. The slice loop is the innermost loop (
not indicated). Each slice is excited and signal-detected prior to any slice being excited a second time for purposes of signal averaging or phase encoding (lines). This loop structure is the most common. (b) Sequential slice looping. All information for a given slice is acquired prior to any excitation for a different slice. In this figure,
. (c) Long-term averaging. All lines for all slices are acquired before performing signal averaging.
Figure 4.11 Three-dimensional slice loop structures. A fourth loop, the partitions loop, is added. Two slice loop structures are commonly used: (a) partitions-in-lines: the partitions gradient amplitudes are varied most frequently; (b) lines-in-partitions: the lines gradient amplitudes are varied most frequently.
Chapter 5: Principles of magnetic resonance imaging – 2
Figure 5.1 Nonselective or rectangular RF pulse: (a) time-domain waveform; (b) frequency-domain waveform.
Figure 5.2 A series of sine waves of different frequencies all in phase at one point in time sum to give an approximation (dashed dark line) of an infinite sinc function (solid dark line).
Figure 5.3 Truncated sinc RF pulse: (a) time-domain waveform; (b) frequency-domain waveform.
Figure 5.4 Gaussian RF pulse: (a) time-domain waveform; (b) frequency-domain waveform.
Figure 5.5 Hyperbolic secant or adiabatic RF pulse: (a) time-domain waveform; (b) frequency-domain waveform.
Figure 5.6 Composite pulses. (a) A 1331 composite pulse is shown with a total excitation angle of 90°. A rotating frame corresponding to the on-resonant frequency is assumed. Prior to the first RF pulse (1), both on-resonance (solid arrow) and off-resonance (dashed arrow) protons are unexcited. At the end of the first RF pulse (2), both will be excited 11.25°. Because of the difference in resonant frequencies, the off-resonance protons become out of phase. The time for the second RF pulse (3) is chosen so that the off-resonance protons are exactly 180°out of phase. At the end of the second RF pulse (4), the on-resonance protons are excited 45°while the on-resonance protons are excited 22.5°. The delay between the second and third RF pulses is chosen so that the off-resonance protons are 180°out of phase with the on-resonance proton (5). A similar delay is chosen between the third and fourth RF pulses (6,7). At the end of the fourth RF pulse (8), the on-resonance protons are rotated 90°(excited), while the off-resonance protons are at 0°(unexcited). (b) A 1331 composite pulse with a total excitation angle of 90°. The end result differs from (a) in that the on-resonance spins are not excited while the off-resonance spins are rotated 90°.
Figure 5.7 Raw data, real (a) and imaginary (b). The raw data matrix has dimensions of
. Each row is a measured signal at a particular
. The number of rows corresponds to
. Signals acquired with high negative amplitude
are displayed at the top, low amplitude
in the middle, and high positive amplitude
at the bottom of the matrix. Each column corresponds to a data point sampled at a different time following the excitation pulse.
Figure 5.8 Image data, magnitude (a) and phase (b). The image data is obtained by performing a two-dimensional Fourier transformation on the data set displayed in Figure 5.7. The rows and columns correspond to the phase-encoding and readout directions. The specification of rows and columns as readout and phase-encoding directions is usually set by the operator.
Figure 5.9 Raw data and corresponding images. (a) Same raw data set as Figure 5.7 except that only the central
data points are kept and zero values are defined for the remaining data. Imaginary portion not shown. (b) Magnitude image data of (a). The image intensity is approximately the same as that of Figure 5.8a, but there is a loss of edge definition, exhibited as blurring in the phantom. (c) Same raw data set as Figure 5.7 except that the central
data points are eliminated, corresponding to 1.56% of the total data set. Imaginary portion not shown. (d) Magnitude image data of (c). The image intensity (central portion of the phantom) is virtually absent, while the edges of the phantom are present.
Figure 5.10 Sequential k-space filling. Each line of k-space (phase-encoding line) corresponds to a measured MR signal. Lines of k-space are acquired serially in time with the data from the maximum negative
acquired first and the data from the maximum positive
acquired last. The center of
k
-space
is acquired halfway through the data collection period. This is the traditional data collection method.
Figure 5.11 Raw data and corresponding image. (a) Raw data set for 192 lines measured with the same
as for Figure 5.7, and zero values are defined for the remaining data. Imaginary portion not shown. (b) Magnitude image data of (a). The FOV is identical and the image intensity is almost identical to that of Figure 5.8a, but there is a slight loss of edge definition, exhibited as blurring in the phantom.
Figure 5.12 Raw data and corresponding image. (a) Raw data set for 192 lines measured with an increased
compared to Figure 5.7, and zero values are defined for the remaining data. Imaginary portion not shown. (b) Magnitude image data of (a). The FOV is reduced compared to that of Figure 5.8a, causing aliasing artifact. (a) (b) GPE
Figure 5.13 Raw data and corresponding image. (a) Same raw data set as Figure 5.7 except that only the upper 136 lines are kept, and zero values are defined for the remaining data. Imaginary portion not shown. (b) Magnitude image data of (a). The spatial resolution and signal intensity are identical to those of Figure 5.8a, but there is a slight increase in noise.
Figure 5.14 Centric
k
-space filling. Lines of
k
-space are acquired serially beginning at the center of
k
-space, then in increasing
amplitudes of alternating polarity.
Figure 5.15 Segmented
k
-space filling. Lines of
k
-space are acquired in groups or segments. The example here shows five segments. One line of data is acquired from each segment before a second line is acquired from any segment. The center of
k
-space is acquired at a time during the scan that is dependent on the number of lines per segment, number of segments, and the order of acquisition within a segment and between segments.
Figure 5.16 Spiral filling of
k
-space, three-dimensional. The
and
amplitudes are varied in an increasing spiral pattern. The lowest-amplitude values are used to acquire the earliest echoes in the scan.
Figure 5.17 Ramped sampling of signal. The dashed lines represent the time points when signal sampling occurs. The
gradient amplitude is not constant during signal sampling. As a result,
, which is proportional to the rectangular gray area, is not constant for all sample points.
Figure 5.18 Radial filling of
k
-space. Each signal is sampled with a
gradient consisting of the sum of two physical gradients with variable amplitudes. The direction of each
varies depending on the particular gradient amplitudes.
Figure 5.19 Spiral filling of
k
-space, two-dimensional. The signal is sampled with the gradients perpendicular to
varying in a spiral fashion.
Figure 5.20 Phased-array coils. A large area such as the spine can be scanned using an array of smaller coils. The images from each coil can be combined to form the final image.
Figure 5.21 Images from the individual coils of an array (outer ring of images) can be combined to produce a single image (central image). Each outer image is sensitive to signals in its vicinity.
Figure 5.22 Principle of parallel imaging. A raw data set is acquired from each coil element. The data sets have a reduced number of lines, allowing for shorter measurement times. For
k
-space-based methods (solid lines), the raw data sets are combined to form a complete raw data set. The complete raw data set is processed to produce the final image. For image-based methods (dashed lines), the raw data sets are processed to produce images, and the images are combined to form the final image.
Figure 5.23 The direction of
should be aligned with the coil geometry axis. The different position of each coil means that the magnetic field will differ at each location.
Figure 5.24 Three-dimensional transverse
T1
-weighted spoiled gradient echo image of abdomen with fat suppression, acquired using a 32-channel body array receiver coil. Scan parameters: GRAPPA with acceleration factors of 2 each in both
and
directions;
TR
, 4.9 ms;
TE
, 2.4 ms; excitation angle, 10°; acquisition matrix,
, 131 and
, 256 with twofold readout oversampling; FOV,
; effective slice thickness, 3.5 mm; time of acquisition, 8 s.
Chapter 6: Pulse sequences
Figure 6.1 Simple timing diagram. The horizontal axis is time during sequence execution. Control signals for three hardware components are illustrated: the RF transmitter and two gradients (labeled RF Pulse, Gradient Pulse, and Gradient Table). The gradient pulse signals are executed the same way for each measurement, as indicated by the trapezoidal shape for the two gradient pulse waveforms. The gradient table signals change amplitude with each measurement and are represented by the multiple horizontal lines.
Figure 6.2 Standard single-echo spin echo sequence timing diagram. These sequences are characterized by a single 180° refocusing pulse, a single detected echo, and a single phase-encoding table. The
TE
time is measured from the middle of the excitation pulse to the center of the echo.
Figure 6.3 Standard multiecho spin echo sequence timing diagram. Two echoes are illustrated. Additional echoes may be generated by adding additional 180° RF pulses, slice-selection gradient pulses, readout gradient pulses, and ADC sampling period. Note the single phase-encoding gradient table. Both
TE
times are measured from the middle of the excitation pulse to the center of the respective echo.
Figure 6.4 Echo train spin echo sequence timing diagram. A three-segment (echo train length of 3) version is illustrated, each with
values of
per segment to acquire
total lines of raw data. The arrows beside the
tables indicate the direction in which the phase-encoding values change from one excitation pulse to the next. Gradient tables on opposite sides of the ADC sampling period have equal amplitudes but opposite polarity. The effective
TE
is the
TE
time (1, 2, or 3) during which the
lines of data are acquired.
Figure 6.5 Spoiled gradient echo sequence timing diagram, two-dimensional method. Because there is no 180° RF pulse, the polarity of the
dephasing gradient pulse (a) is opposite that of the readout gradient pulse applied during signal detection. Gradient spoiling is illustrated at the end of the kernel. The
TE
time is measured from the middle of the excitation pulse to the center of the echo.
Figure 6.6 A series of equally spaced RF pulses produces spin echoes that form at the time of the subsequent RF excitation pulses. Once a steady state is reached (following several
time periods), signal is induced prior to and following the excitation pulse. The preexcitation pulse signal
is strictly echo reformation. The postexcitation pulse signal
is a combination of echo and free induction decay. Images can be produced from either signal or from both.
Figure 6.7 Refocused gradient echo sequence timing diagrams. Instead of spoiling the transverse magnetization, rephasing gradients are used to maintain the transverse magnetization as much as possible. (a) Postexcitation imaging sequence; (b) preexcitation imaging sequence.
Figure 6.8 Coronal image acquired with fully refocused gradient echo technique. The arrow indicates a signal cancellation artifact due to interference between the pre-and postexcitation signals. (Reproduced with permission of H. Cecil Charles, Duke University.)
Figure 6.9 Echo planar imaging sequence timing diagrams. A spin echo excitation scheme and an echo train length of 8 are illustrated. (a) Constant phase encoding. The phase-encoding gradient is on for the entire data collection period. Each data point of each echo has a unique amount of
influence. (b) Blipped phase encoding. The phase-encoding gradient is applied incrementally prior to detection of each echo. Each echo is influenced by a different amount of
, but each data point within the echo has the same amount of
influence.
Figure 6.10 Standard inversion recovery sequence timing diagram. The
TI
time is measured from the center of the inversion (initial 180°) pulse to the center of the excitation pulse. The
TE
time is measured from the center of the excitation pulse to the center of the echo.
Figure 6.11 Echo train inversion recovery sequence timing diagram. A three-segment (echo train length of 3) version is illustrated, each segment with
values for
. The
TI
time is measured from the center of the inversion (initial 180°) pulse to the excitation pulse. The effective
TE
is the
TE
time (1, 2, or 3) during which the
lines of data are acquired.
Figure 6.12 Inversion recovery looping modes. (a) Inversion for a given slice is followed immediately by echo acquisition. This is typically used when the
TI
time is relatively short. (b) All inversion pulses for all slices are performed followed by echo acquisition. This is typically used when the
TI
time is relatively long.
Figure 6.13
T1
recovery curves for inversion recovery sequences. The 180° inversion pulse inverts the net magnetization for all tissues. Tissue with a short
T1
time (solid curve) recovers faster than tissue with a long
T1
(dashed curve) time. For a
TI
time of (a), both tissues contribute significant negative amplitude signal. For a
TI
time of (b), the short
T1
tissue contributes positive amplitude signal while the long
T1
tissue contributes negative amplitude signal. For a
TI
time of (c), the short
T1
tissue contributes significant positive amplitude signal and the long
T1
tissue contributes minimal signal. If the signal polarity is considered (phase-sensitive IR sequences), signal difference will be seen at all three
TI
times. If the signal polarity is ignored (absolute value or magnitude IR sequences), no difference in signal between the two tissues will be seen at time (b).
Figure 6.14 Magnitude (a) and phase-sensitive (b) inversion recovery images. All other measurement parameters are equal: pulse sequence, echo train spin echo, five echoes;
TR
, 7000 ms;
TE
, 14 ms;
TI
, 140 ms; acquisition matrix;
, 224 and
, 256;
FOV
,
RO;
, 1; slice thickness, 5 mm.
Figure 6.15 Two-dimensional MP sequence timing diagram,
T1
-weighted. A single 180° inversion RF pulse is applied once per scan (section A). This inversion provides significant
T1
weighting to
. The portion of the sequence indicated by
is repeated for the
and
desired.
Figure 6.16 Two-dimensional MP sequence timing diagram,
T2
-weighted. A 90°– 180°– 90° RF pulse train is applied once prior to the data collection scheme (section A). The first two pulses produce
T2
weighting to
, which is restored to the longitudinal direction by the final pulse prior to the data collection. The portion of the sequence indicated by
is repeated for the
and
desired.
Figure 6.17 Three-dimensional MP sequence timing diagram,
T1
-weighted. A 180° inversion RF pulse is applied followed by encoding in the slice
direction (section A). This inversion provides significant
T1
weighting to
. The portion of the sequence indicated by
is repeated for the
desired. The entire process (A and B) is repeated for the
and
desired.
Chapter 7: Measurement parameters and image contrast
Figure 7.1
TR
effects on image contrast. Longer
TR
allows more time for
T1
relaxation and produces more signal from tissues with long
T1
values. Other measurement parameters are: pulse sequence, spin echo;
TE
, 30 ms; acquisition matrix,
, 224 and
, 256; FOV,
RO;
, 1; slice thickness, 5 mm. (a)
TR
of 500 ms; (b)
TR
of 2000 ms. Note reversal of contrast between gray matter and white matter in (b) compared to (a).
Figure 7.2
TE
effects on image contrast. Longer
TE
allows more time for
T2
relaxation and produces more signal from tissues with long
T2
values. Other measurement parameters are: pulse sequence, spin echo;
TR
, 2000 ms; acquisition matrix,
, 224 and
, 256; FOV,
;
, 1; slice thickness, 5 mm. (a)
TE
of 30 ms; (b)
TE
of 80 ms. Note bright signal from cerebrospinal fluid (CSF) in (b) compared to (a).
Figure 7.3
TI
effects on image contrast. Longer
TI
allows more time for
T1
relaxation following the inversion pulse. The choice of
TI
can cause signal suppression of different tissues. Other measurement parameters are: pulse sequence, echo train spin echo, five echoes;
TR
, 7000 ms;
TE
, 14 ms; acquisition matrix,
, 224 and
, 256; FOV,
;
, 1; slice thickness, 5 mm. (a)
(fat suppression); (b)
(CSF suppression).
Figure 7.4 Excitation order and crosstalk. If the slices are closely spaced, the bases of adjacent slices overlap (crosshatched regions). Tissues located in this overlap region experience RF pulses from both slices and become saturated. This double excitation called crosstalk, causes reduced signal from these regions. The order of slice excitation also determines the contribution of crosstalk to the image intensity. Ascending order of excitation (second row) acquires data from adjacent slice positions in successive time periods. Interleaved ordering (third row) acquires data from every other slice position first, then acquires data from the intermediate positions. This ordering will minimize the effects of crosstalk between all slices.
Chapter 8: Signal suppression techniques
Figure 8.1 Standard single-echo spin echo sequence timing diagram including a spatial presaturation pulse. The saturation pulse (labeled SAT) is applied prior to the primary slice excitation pulse (labeled 90°). The RF pulse center frequency and bandwidth and the gradient amplitudes for the presaturation pulse are independent of these variables for the slice excitation pulses.
Figure 8.2 Application of a spatial presaturation pulse to moving tissue will suppress signal from that tissue. Measurement parameters are: pulse sequence, spin echo;
TR
, 500 ms;
TE
, 16 ms; excitation angle, 90° ; acquisition matrix,
N
PE
, 192 and
N
RO
, 256; FOV,
RO;
N
SA
, 3; slice thickness, 4 mm. (a) No presaturation pulse; (b) coronal spatial presaturation pulse, suppressing artifact from swallowing.
Figure 8.3 Standard single-echo spin echo sequence timing diagram, including a frequency- selective presaturation pulse. The saturation pulse (labeled SAT) is applied prior to the primary slice excitation pulse (labeled 90°). The RF pulse center frequency and bandwidth for the presaturation pulse are independent of these variables for the slice excitation pulses. Note the absence of the associated gradient pulse for the presaturation pulse compared to Figure 8.1.
Figure 8.4 Magnetization transfer suppression. Mobile or “free” tissue water has protons with long
T2
values and produces a narrow resonance peak. Water adsorbed or “bound” to macromolecules has protons with short
T2
values and produces a wide resonance peak normally not visualized in an image. Both types of water protons have the same resonant frequency. The magnetization transfer RF pulse is applied at a frequency different (off-resonance) from the water to saturate the bound water protons. Exchange between the bound and free water transfers the saturation to the free water protons, reducing signal intensity from the free water.
Figure 8.5 Effects of magnetization transfer in three-dimensional MR angiography. Application of MT pulse suppressed background signal from gray and white matter, enabling better visualization of blood vessels. An apparent increase in signal from suborbital fat is observed (arrows). Measurement parameters are: pulse sequence, three-dimensional refocused gradient echo, postexcitation;
TR
, 42 ms;
TE
, 7 ms; excitation angle, 25°; acquisition matrix,
N
PE
, 192 and
N
RO
512 with twofold readout oversampling; FOV,
RO;
N
SA
, 1; effective slice thickness, 0.78 mm. (a) No MT pulse; (b) MT pulse.
Figure 8.6 Effects of magnetization transfer in
T1
-weighted imaging following contrast administration. Application of MT pulse suppresses background signal from normal matter, enabling better visualization of contrast-enhanced tissues such as tumors or vascular structures. (a) No MT pulse; (b) MT pulse.
Figure 8.7 Frequency spectrum of fat and water. Fat saturation applies an additional RF excitation pulse centered at the fat resonant frequency. This pulse is applied prior to the primary slice excitation pulse, so that the signal from the slice is produced primarily from the water.
Figure 8.8 Frequency-selective saturation (fatsat) pulse is applied to suppress signal from fat protons. (a) With a homogeneous magnetic field, the suppression of fat is uniform throughout the slice; (b) with a nonhomogeneous magnetic field, the saturation pulse suppresses fat well in one region of the image and poorly in another region (arrows).
Figure 8.9 Water excitation image acquired using a composite RF pulse. Measurement parameters:pulse sequence, three-dimensional gradient echo, combination pre-and postexcitation;
TR
, 25.4 ms;
TE
, 9 ms; excitation angle, 35° cumulative; acquisition matrix,
N
PE
, 256 and
N
RO
, 256; FOV,
RO;
N
SA
, 1; effective slice thickness, 1.56 mm.
Figure 8.10 Dixon method for fat suppression: (a) water image; (b) fat image. Poor magnet homogeneity results in a phase wrap, causing the fat and water to appear in the incorrect images (arrows).
Chapter 9: Artifacts
Figure 9.1 Flow misregistration artifact, through-plane. (a) Flow fast compared to measurement technique produces “zipper”-like artifact (arrow). Measurement technique: pulse sequence, spin echo;
TR
,
;
TE
,
; acquisition matrix,
, 192 and
; FOV, 172 mm PE × 230 mm RO;
, 1; slice thickness, 5 mm. PE direction: R–L; RO direction: A–P. (b) Periodic flow from the aorta will be misregistered as multiple ghosts (arrows). Measurement parameters are: pulse sequence, two-dimensional spoiled gradient echo;
TR
,
;
TE
,
; excitation angle, 70°; acquisition matrix,
, 134 and
; FOV, 262 mm
;
, 1; slice thickness, 5 mm; PE direction, A–P; RO direction, L–R.
Figure 9.2 Flow misregistration artifact, in-plane. (a) In-plane blood flow will produce severe signal misregistration (arrow). Measurement technique: pulse sequence, spin echo;
TR
,
;
TE
,
; acquisition matrix,
, 192 and
, 256; FOV,
;
, 1; slice thickness, 5 mm; PE direction: R–L; RO direction: A–P. (b) Flow misregistration artifact. Flowing CSF will be misregistered as a ghost canal (arrow). Measurement parameters are: pulse sequence, spin echo;
TR
,
;
TE
, 90 ms; excitation angle, 90°; acquisition matrix,
, 192 and
, 256 with twofold frequency oversampling; FOV,
;
, 1; slice thickness, 5 mm; PE direction,
; RO direction,
.
Figure 9.3 Respiratory motion artifact. Extraneous ghost images are generated due to motion of the abdominal wall during data acquisition (arrows). The number and severity of the ghosts depends on the
TR
, respiration rate, and the particular measurement technique. (a) Pulse sequence, two-dimensional spoiled gradient echo;
TR
, 159 ms; (b) pulse sequence, echo train spin echo;
TR
, 4000 ms.
Figure 9.4 Effects of oversampling. (a) Without frequency oversampling, frequencies for the protons within the arms exceed the Nyquist limit and are aliased or incorrectly mapped into the image (arrows). Measurement parameters are: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 140 ms;
TE
, 4 ms; excitation angle, 80°; acquisition matrix,
, 128 and
, 256; FOV,
;
, 1; slice thickness, 8 mm. (b) Same as (a) except with frequency oversampling. The frequencies for the protons within the arms are measured accurately by increasing the number of readout data points measured during the same sampling time while maintaining the same
. Only frequencies corresponding to the FOV selected are stored so that the arms are excluded from the final image. Measurement parameters are the same as (a), except that
.
Figure 9.5 Three-dimensional volume scan showing high-frequency aliasing in the slice-selection direction. The superior part of the brain (arrow) appears in the inferior slices of the imaging volume.
Figure 9.6 Small FOV aliasing artifacts with parallel acquisitions: (a) image-based parallel acquisition; (b)
k
-space-based acquisition. Note the aliasing artifact in the lung field in (a) (arrow) due to the anatomical region being larger than the scan FOV. This area is artifact-free in (b).
Figure 9.7 Chemical shift artifact. (a) Note alternate bands of light and dark at the interface between the vertebrae and disk (arrows). Measurement parameters: pulse sequence, spin echo; Receiver bandwidth, 20 kHz; readout direction,
. (b) A complete misregistration of fat from the bone marrow of the skull (arrow). Measurement parameters are: pulse sequence, spin echo EPI; phase-encoding direction,
.
Figure 9.8 Phase cancellation artifact. Other measurement parameters are: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 164 ms; excitation angle, 70°; acquisition matrix,
, 134 and
, 256; FOV,
;
, 1; slice thickness, 9 mm. (a) In-phase image (
TE
, 4.5 ms). Fat and water protons have the same phase and contribute in the same fashion to the image contrast. (b) Out-of-phase image (
TE
, 2.2 ms). Fat and water protons have opposite phases and contribute in opposite fashion to the image contrast. For voxels with equal amounts of fat and water, such as at the interface between liver or kidney and retroperitoneal fat, cancellation of signal occurs, producing a dark band (arrows).
Figure 9.9 Truncation artifact. Image acquisition with asymmetric sampling produces a banding artifact (arrow), originating from the high signal of subcutaneous fat. Measurement parameters: pulse sequence, two-dimensional spoiled gradient echo;
TR
,
;
TE
,
; excitation angle, 80°; acquisition matrix,
, 144 and
, 256 with twofold oversampling; FOV, 262 mm PE × 350 mm RO;
, 1.
Figure 9.10 Symmetric versus asymmetric echo sampling. (a) In symmetric sampling, the echo is maximum (
TE
) midway through the sampling period, so that both sides of the echo signal are measured equally. Filtering of the data can be performed in a symmetrical fashion. (b) In asymmetric sampling, the echo maximum is in the early portion of the sampling period. Significant signal is present when the sampling begins and filtering of the signal is difficult.
Figure 9.11 Raw data filtering. All vertical and horizontal scales are maintained in all figures. (a) Unfiltered sinc pulse waveform, time domain (left) and magnitude frequency domain (right) (note sawtooth pattern at top of right figure); (b) Gaussian filter (left), Gaussian-filtered sinc pulse waveform, time domain (center) and magnitude frequency domain (right). (note that sawtooth pattern at top of right figure is reduced but still present); (c) Fermi filter (left), Fermi-filtered sinc pulse waveform, time domain (center) and magnitude frequency domain (right) (note rounded top and large width at base of right figure); (d) Hanning filter (left), Hanning-filtered sinc pulse waveform, time domain (center) and magnitude frequency domain (right).
Figure 9.12 Insufficient spoiling, raw data and image. Due to the imperfect nature of RF pulses, additional signal can be produced by spins at the edge of the slice. If the spurious signal is generated by the 180° pulse in a spin echo pulse sequence, it is not affected by the phase encoding process and appears as a constant signal in
k
-space (arrow in a). The resulting image shows a bright line near the edge of the image FOV (arrow in b).
Figure 9.13 Echo timing plots. (a) Time
t
1
is less than
, typical of a standard short
TE
/ long
TE
spin echo pulse sequence. Five echoes are formed. Two are used in routine imaging (echoes 1 and 3). The stimulated echo occurs at time
(echo 2). (b) Time
t
1
is greater than
, typical of a spin echo pulse sequence with a short
TE
and a spatial presaturation pulse. Four echoes are formed. Two are used in routine imaging (echoes 1 and 2). The stimulated echo occurs at time
(echo 3).
Figure 9.14 Signal contributions from undesired echoes produce banding artifacts (arrow). These echoes can be minimized through the use of coherence spoiling, either gradient-or RF-based.
Figure 9.15 RF spoiling. (a) No spoiling. The phase of the RF excitation pulse in a rotating frame is the same for all RF pulses. The net magnetization will be rotated along the same axis following each RF excitation pulse. (b) RF spoiling. The phase of the RF excitation pulse in a rotating frame is incremented so that each RF pulse has a different phase. The net magnetization will be rotated along different axes following each RF excitation pulse.
Figure 9.16 Magnetic susceptibility difference artifact. An increase in
TE
causes increased sensitivity to
T2
* in gradient echo pulse sequences. Increased distortions occur in areas where there are significant differences in magnetic susceptibility, such as in the posterior fossa (arrows). Other measurement parameters: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 170 ms; excitation angle, 30°; acquisition matrix,
, 224 and
, 256; FOV, 201 mm PE × 230 mm RO;
, 1. (a)
TE
, 4 ms; (b)
TE
, 15 ms.
Figure 9.17 Magnetic susceptibility difference artifact. As paramagnetic contrast agents accumulate in the kidneys, the local magnetic field is distorted, causing enhanced dephasing to the protons in the vicinity and produces signal voids (arrows). Measurement parameters are: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 140 ms;
TE
, 4 ms; acquisition matrix,
, 128 and
, 256 with twofold oversampling; FOV,
;
,1.
Figure 9.18 Radial artifact. Insufficient number of sample planes in a radial scan produces a scatter or “star” artifact. The primary scatter lines are indicated by the white arrows.
Figure 9.19 Magnetic susceptibility difference artifact. Surgical clips produce a void of signal caused due to the significant magnetic field distortion. (a) Spoiled gradient echo sequence shows significant signal distortion (arrow). (b) Single-shot echo train spin echo at the same level shows less signal loss (arrow). (c) Single-shot echo train spin echo with fat saturation shows incomplete saturation due to field distortions (arrows).
Figure 9.20 Stainless-steel aneurysm clip produces severe field distortion. Pulse sequence: spin echo.
Figure 9.21 Spikes. Transient electrical discharges (spikes) during the data collection period produce a banding pattern that is superimposed across the entire imaging field. The direction and spacing of the bands depend on the timing of the discharge relative to the collection of the central phase-encoding steps.
Figure 9.22 External interference: (a) artifact due to electrical source outside the scan room; (b) interference from portable patient medication unit operating in scan room during the measurement (arrows).
Chapter 10: Motion artifact reduction techniques
Figure 10.1 The direction of motion artifacts is determined by the phase-encoding gradient. Blood flow during the scan produces motion artifact (arrow). (a) RO: left/right, PE: anterior/ posterior; (b) RO: anterior/posterior, PE: left/right.
Figure 10.2 Single-shot echo train spin echo image of abdomen. Note lack of motion artifact from heart and bowel.
Figure 10.3 Triggered data collection. The R wave is used as a timing signal. A trigger delay (TD) can be used to initiate data collection at any time desired during the cardiac cycle. (a) Nonsegmented measurement. For multislice single-phase imaging, one
line for each slice is acquired per heartbeat. Information for a particular slice is acquired at the same time following the R wave. The
T1
contrast is based on the R–R time interval rather than the user-defined
TR
. Alternatively, a multislice, multiphase acquisition may be carried out in which slices are acquired at different times during the cardiac cycle at the same position. (b) Segmented measurement. For single-slice imaging, multiple
lines are acquired for a particular slice following the R wave. For multislice imaging, multiple
lines are acquired for multiple slices following the R wave.
Figure 10.4 Short-axis
T1
-weighted image acquired using a triggered multislice mode.
Figure 10.5 Cine heart images acquired at different phases of cardiac cycle. Measurement parameters: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 35.4 ms;
TE
, 1.5 ms; excitation angle, 65°; acquisition matrix,
, 108 and
, 128; FOV, 169 mm PE × 200 mm RO;
, 1. (a) 92 ms following R wave; (b) 552 ms following R wave.
Figure 10.6 Flow compensation. Use of flow compensation gradient pulses will map moving protons such as cerebrospinal fluid to their proper location. Measurement parameters are: pulse sequence, spin echo;
TR
, 2500 ms;
TE
, 90 ms; excitation angle, 90°; acquisition matrix,
, 192 and
, 256 with twofold readout oversampling; FOV,
;
, 1; slice thickness, 5 mm. (a) No flow compensation. Misregistration artifact from CSF flow appears anterior to the spinal canal (arrow). (b) First-order flow compensation in readout and slice-selection directions. CSF is properly mapped into the spinal canal.
Figure 10.7 Radial motion compensation. Each scan measures a portion of
k
-space (dark lines). Some of the measured lines contain points acquired at the center of
k
-space.
Chapter 11: Magnetic resonance angiography
Figure 11.1 Flow selection in MR angiography. (a) In the absence of a spatial presaturation pulse, flowing blood entering the imaging volume is visualized regardless of the direction of the flow. (b) A spatial presaturation pulse that saturates flowing blood prior to its entry into the imaging volume suppresses signal from that volume.
Figure 11.2 Time-of-flight effect. During data collection, the imaging volume experiences multiple RF pulses. Flowing blood (colored box) experiences the first RF pulse (upper left). During the first
TR
period, the excited blood volume moves (upper right) and only a portion experiences the second RF pulse. During the second
TR
time period, the initial volume of blood continues to move (lower left). By the end of a third
TR
time period, the initial volume of blood is entirely outside the volume of excitation and does not contribute to the detected signal (lower right).
Figure 11.3 Time-of-flight MRA showing effects of magnetization transfer pulse. Other measurement parameters are: pulse sequence, three-dimensional refocused gradient echo, post excitation;
TR
, 42 ms;
TE
, 7 ms; excitation angle, 25°; acquisition matrix,
, 192 and
, 512 with twofold readout oversampling; FOV,
;
, 1; effective slice thickness, 0.78 mm. (a) Source image, one from data set acquired without MT pulse; (b) transverse post processed image of (a); (c) source image, one from data set acquired with MT pulse; (d) transverse post processed image of (c). Note reduction of background gray and white matter in (c) compared to (a) and improved visualization of vessels in (d) compared to (b) (arrow).
Figure 11.4 TONE RF pulse. (a) Standard excitation pulses provide uniform energy deposition across the slice, which gradually increases saturation of, and reduces the transverse magnetization from, blood located at the exit side of the slice, causing a loss of signal. (b) Nonuniform excitation pulses known as ramped or tilted optimized nonuniform excitation (TONE) RF pulses increase the excitation across the slice. Although the amount of saturation increases, the resulting transverse magnetization remains constant, so that the blood signal remains uniform throughout the imaging volume.
Figure 11.5 Time-of-flight MRA showing effects of TONE RF pulse. Measurement parameters are: pulse sequence, three-dimensional refocused gradient echo, post excitation;
TR
, 42 ms;
TE
, 7 ms; excitation angle, 25°; acquisition matrix,
, 192 and
, 512 with twofold readout oversampling; FOV,
;
, 1; effective slice thickness, 0.78 mm. (a) Sagittal projection of volume using normal uniform excitation pulse; (b) sagittal projection of volume using TONE excitation pulse. Note improved signal from vessels (arrow).
Figure 11.6 MR angiography of aorta and renal arteries following bolus administration of gadolinium–chelate contrast agent. Measurement parameters are: pulse sequence, three-dimensional refocused gradient echo, postexcitation;
TR
, 3 ms;
TE
, 1.1 ms; excitation angle, 20°; acquisition matrix,
, 128 and
, 256; FOV,
;
, 1; effective slice thickness, 2.0 mm. (a) Source image, one from data set. Image was obtained following subtraction of unenhanced scan from enhanced scan. (b) Coronal maximum intensity projection image of (a).
Figure 11.7 Peripheral MRA, postprocessed images. Images were acquired at three different scan table positions. Measurement parameters: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 300 ms:
TE
, 3.6 ms; excitation angle, 50°; acquisition matrix,
, 144 and
, 256; FOV,
;
, 1; slice thickness, 4.0 mm; pulse triggered. (Reproduced with permission of H. Cecil Charles, Duke University.)
Figure 11.8 Phase-contrast MRA. Venc, 30 cm s
−1
in all directions. Other measurement parameters: pulse sequence, two-dimensional spoiled gradient echo;
TR
, 211 ms;
TE
, 8.4 ms; excitation angle, 15°; acquisition matrix,
, 192 and
, 256; FOV, 240 mm PE × 240 mm RO;
, 1; slice thickness, 40.0 mm. (a) Flow image, flow direction A–P; (b) flow image, flow direction R–L; (c) flow image, flow direction S–I; (d) magnitude sum of images (a) to (c); (e) phase image, flow direction A–P; (f) phase image, flow direction R–L; (g) phase image, flow direction S–I.
Figure 11.9 Maximum intensity projection (MIP). The MR images are acquired so that moving blood has pixels of maximal intensity. The MIP process maps the pixels of maximum intensity into a single projection or view, regardless of which slice the pixel was located in. Changing the direction of projection provides a different perspective of the vessels.
Figure 11.10 Maximum intensity projection. Note lack of vessel in (a) (arrow), which is caused by improper exclusion (b). Processing of the complete data set shows the entire vessel (c).
Chapter 12: Advanced imaging applications
Figure 12.1 Spin echo pulse sequence showing diffusion gradients, known as the Stejskal – Tanner approach.
G
is the amplitude for each of the gradient pulses,
t
is the duration of the gradient pulse during which the diffusion weighting occurs, and
T
is the time between the two pulses.
Figure 12.2 Diffusion-weighted EPI sequence: (a)
; (b)
; (c)
. Normal tissue has moderate diffusion of water, while tissue under stress, such as that at risk for a stroke, has restricted motion of tissue water and shows increased signal on image with significant diffusion sensitivity. (d) ADC map, calculated from images (a), (b), and (c). Low pixel amplitudes indicate restricted water movement. High pixel amplitudes indicate free water movement.
Figure 12.3 Ellipsoids representing principal axes for diffusion. (a) Isotropic diffusion: diffusion is equally likely in all directions and is represented by a sphere. (b) Anisotropic diffusion: diffusion is more likely in one or more directions. The direction that is most likely has the largest vector length.
Figure 12.4
T1
-weighted two-dimensional spoiled gradient echo imaging of liver following administration of gadolinium–chelate contrast agent. (a) Image acquired prior to contrast administration. (b) Image acquired immediately following administration. Contrast agent is in the hepatic arterial phase, as evidenced by the nonopacified hepatic vein (arrow). (c) Image acquired 45 seconds following administration. Contrast agent is now in the capillary phase, as evidenced by its presence in the hepatic vein (arrow).
Figure 12.5
T1
-weighted three-dimensional volume spoiled gradient echo imaging of breast following administration of gadolinium – chelate contrast agent. (a) Precontrast image, lacking evidence of lesion; (b – d) serial images acquired every 48 seconds following contrast agent administration. Note increased signal from lesion in later images.
Figure 12.6
image of normal lung acquired following inhalation of hyperpolarized helium gas. Note significant signal in trachea and upper lobes of lungs and lack of signal from other tissue in the body. Measurement parameters: pulse sequence, two-dimensional refocused gradient echo, postexcitation;
TR
, 25 ms;
TE
, 10 ms; acquisition matrix,
, 128 and
, 256; FOV,
. (Reproduced with permission of James R. MacFall, Duke University.)
Chapter 13: Magnetic resonance spectroscopy
Figure 13.1 Typical
spectrum from normal brain. Measurement parameters: pulse sequence, two-dimensional CSI PRESS;
TR
, 1500 ms;
TE
, 144 ms;
, 4; voxel size,
.
Figure 13.2 Chemical structures of creatine (a) and phosphocreatine (b). Because the phosphate group is located several bonds away, it has very little influence on the molecular environment of the methyl hydrogens (boldface type). As a result, the signals for the methyl hydrogens in both molecules have the same frequency.
Figure 13.3 PRESS
spectra of lactate (doublet) and acetate (singlet). Choice of
TE
affects the relative polarity of lactate peak compared to acetate, due to modulation of spin-coupled
atoms. (a)
; (b)
.
Figure 13.4 PRESS pulse sequence timing diagram. The CHESS RF pulse is used for suppression of water.
Figure 13.5 STEAM pulse sequence timing diagram. The CHESS RF pulse is used for suppression of water.
Figure 13.6 Volume-selective pulse sequence timing diagrams: (a) two-dimensional; (b) three-dimensional. The CHESS RF pulse is used for suppression of water.
Figure 13.7 Typical spectroscopy postprocessing steps.
Figure 13.8 Phase-corrected complex spectrum resulting from Fourier transformation of FID signal. (a) Real portion of spectrum, also known as the absorption spectrum; (b) imaginary portion of spectrum, also known as the dispersion spectrum.
Figure 13.9
1
H two-dimensional CSI PRESS spectra from different voxel locations. Note the difference in spectral patterns.
Figure 13.10 Normal prostate spectra.
Figure 13.11
two-dimensional CSI PRESS spectra acquired at 1.5 T (a) and 3.0 T (b). The scan parameters are identical except that spectrum (a) had four averages, whereas spectrum (b) had one average.
Chapter 14: Instrumentation
Figure 14.1 Block diagram of an MRI system.
Figure 14.2 The effective gradient
is the vector sum of the gradients
in all three directions.
Figure 14.3 Polarization of RF wave. (a) Linear polarization. A plane wave can be thought of as the sum of two circular waves rotating in opposite directions. Because the nuclear precession is in only one direction, only one component interacts with the precessing spins (solid arrow) and is effective at resonance absorption. The other component will be absorbed as heat. (b) Circular polarization. A plane wave can be divided into two parts. If one part is shifted 90° relative to the other wave, each part will have one component interacting with the precessing spins (solid arrows), which add together. The counter-rotating components (broken arrows) are always 180° from each other and cancel. The resulting wave is a circular or helical wave with no out-of-phase component.
Chapter 15: Contrast agents
Figure 15.1 Exchange of water molecules in coordination sphere of gadolinium – chelate contrast agent. The chelate molecule causes steric hindrance (crowding) around the gadolinium atom, restricting its access. An excited water molecule is small enough to reach the inner sphere and transfer its energy to the gadolinium ion, then leaves the complex unexcited as another molecule replaces it.
Figure 15.2
T1
-weighted spin echo images before and after administration of gadolinium–chelate contrast agent: (a) no contrast agent present; (b) contrast agent present. Signal increase due to presence of agent.
Figure 15.3 MRA of renal arteries following half - dose of gadolinium – chelate contrast agent: (a) acquired image from three - dimensional volume acquisition; (b) MIP projection.
Figure 15.4
T2
-weighted single-shot echo train spin echo image of liver following administration of superparamagnetic iron oxide contrast agent. The
T2
relaxation times of the normal tissue are reduced by the iron oxide, while the metastatic lesion shows less agent uptake.
Chapter 17: Clinical applications
Figure 17.1 Sagittal spin echo
T1
-weighted head image.
TR
,
;
TE
, 7.7 ms.
Figure 17.2 Transverse echo train spin echo
T2
-weighted head image.
TR
,
; effective
TE
,
; echo train length, 5.
Figure 17.3 Transverse echo train spin echo fluid attenuated head image,
TR
, 9000 ms; effective
TE
,
;
TI
,
.
Figure 17.4 Sagittal echo train spin echo
T1
-weighted lumbar spine image.
TR
,
; effective
TE
,
; echo train length, 3. Anterior spatial presaturation pulse is used to suppress peristalsis and respiration artifacts.
Figure 17.5 Sagittal echo train
T2
-weighted lumbar spine image.
TR
,
; effective
TE
,
; echo train length, 17. Anterior spatial presaturation pulse is used to suppress peristalsis and respiration artifacts.
Figure 17.6 Transverse echo train spin echo
T1
-weighted lumbar spine image.
TR
,
; effective
TE
,
; echo train length, 3.
Figure 17.7 Coronal spin echo
T1
-weighted knee image.
TR
,
;
TE
,
.
Figure 17.8 Sagittal echo train spin echo proton density–weighted knee image.
TR
, 3000 ms; effective
TE
,
; echo train length, 6.
Figure 17.9 Transverse echo train spin echo proton density–weighted knee image.
TR
,
; effective
TE
,
; echo train length, 11.
Figure 17.10 Coronal echo train spin echo proton density–weighted knee image with fat saturation.
TR
, 3000 ms; effective
TE
,
; echo train length, 6.
Figure 17.11 Oblique sagittal spin echo
T1
-weighted shoulder image.
TR
,
;
TE
,
.
Figure 17.12 Transverse echo train spin echo proton density–weighted shoulder image with fat saturation.
TR
,
; effective
TE
,
; echo train length, 7.
Figure 17.13 Short axis echo train spin echo dark-blood cardiac image.
TR
, 1142 ms; effective
TE
, 32 ms; echo train length, 21.
Figure 17.14 Four-chamber steady-state gradient echo bright-blood cardiac image.
TR
, 36.1 ms;
TE
, 1.8 ms.
Figure 17.15 Oblique coronal steady-state gradient echo bright-blood heart image.
TR
,
;
TE
,
.
Figure 17.16 Coronal single-shot echo train
T2
-weighted image of liver and kidneys. Effective
TE
,
.
Figure 17.18 Transverse three-dimensional volume
T1
-weighted spoiled gradient echo liver image with fat suppression.
TR
,
;
TE
,
; excitation angle, 10°.
Figure 17.19 Transverse three-dimensional volume spoiled gradient echo
T1
-weighted liver image with fat suppression, following gadolinium – chelate contrast administration.
TR
,
;
TE
,
; excitation angle, 10°.
Figure 17.20 Coronal single-shot echo train
T2
-weighted image of bowel. Effective
TE
,
.
Figure 17.21 Coronal three-dimensional volume spoiled gradient echo
T1
-weighted image of kidneys with fat suppression, following gadolinium – chelate contrast administration.
TR
,
;
TE
,
; excitation angle, 10°.
Figure 17.22 Coronal single-shot echo train spin echo
T2
-weighted image of common bile duct, following fat suppression.
TR
,
; effective
TE
,
;
TI
,
.
Figure 17.23 Transverse echo train spin echo proton density–weighted image of pelvis.
TR
, 4500 ms; effective
TE
,
; echo train length, 7.
Figure 17.24 Transverse echo train spin echo proton density–weighted image of pelvis.
TR
,
; effective
TE
,
; echo train length, 7.
Figure 17.25 Coronal three-dimensional volume spoiled gradient echo image, following gadolinium–chelate contrast administration.
TR
, 4.9 ms;
TE
,
; excitation angle, 10°.
Figure 17.26 Diffusion-weighted imaging of the prostate at 3 T: (a)
b
-value, 50 s mm
−2
; (b)
b
-value, 800 s mm
−2
; (c) apparent diffusion coefficient map. Courtesy of Siemens AG, in cooperation with Dr. Engelhardt, Hospital Martha-Maria, Nuernberg, Germany.
List of Tables
Chapter 1: Production of net magnetization
Table 1.1 Constants for Selected Nuclei of Biological Interest
Chapter 6: Pulse sequences
Table 6.1 Spin echo pulse sequence acronyms
Table 6.2 Gradient echo pulse sequence acronyms
Table 6.3 Inversion recovery pulse sequence acronyms
Chapter 7: Measurement parameters and image contrast
Table 7.1 Measurement effects – extrinsic parameters
Table 7.2 Measurement effects – intrinsic parameters
Chapter 9: Artifacts
Table 9.1 Phase cycling
TE
times
Chapter 15: Contrast agents
Table 15.1 Thermodynamic and kinetic stability measurements of gadolinium chelates (from Port et al., 2008). (Abbreviations:
K
cond
= conditional stability constant at physiological pH;
K
therm
= thermodynamic stability constant;
T
1/2
= half-life.)
Table 15.2 Acute adverse reactions after contrast media injection (from ACR, 2013)
Table 15.3 Chronic kidney disease (CKD) classification based upon the glomerular filtration rate (GFR) (from CKD Work Group, 2013)