MRI in Practice continues to be the number one reference book and study guide for the registry review examination for MRI offered by the American Registry for Radiologic Technologists (ARRT). This latest edition offers in-depth chapters covering all core areas, including: basic principles, image weighting and contrast, spin and gradient echo pulse sequences, spatial encoding, k-space, protocol optimization, artefacts, instrumentation, and MRI safety. * The leading MRI reference book and study guide. * Now with a greater focus on the physics behind MRI. * Offers, for the first time, equations and their explanations and scan tips. * Brand new chapters on MRI equipment, vascular imaging and safety. * Presented in full color, with additional illustrations and high-quality MRI images to aid understanding. * Includes refined, updated and expanded content throughout, along with more learning tips and practical applications. * Features a new glossary. MRI in Practice is an important text for radiographers, technologists, radiology residents, radiologists, and other students and professionals working within imaging, including medical physicists and nurses.
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Library of Congress Cataloging-in-Publication Data
Names: Westbrook, Catherine, author. | Talbot, John (Writer on magnetic
resonance imaging), author.
Title: MRI in practice / by Catherine Westbrook, John Talbot.
Description: Fifth edition. | Hoboken, NJ : Wiley, 2018. | Includes
bibliographical references and index. |
Identifiers: LCCN 2018009382 (print) | LCCN 2018010644 (ebook) | ISBN
9781119391999 (pdf) | ISBN 9781119392002 (epub) | ISBN 9781119391968
Subjects: | MESH: Magnetic Resonance Imaging–methods | Magnetic Resonance
Classification: LCC RC78.7.N83 (ebook) | LCC RC78.7.N83 (print) | NLM WN 185
| DDC 616.07/548–dc23
LC record available at https://lccn.loc.gov/2018009382
Cover image: Courtesy of John Talbot
Cover design by Wiley
About the companion website
1 Basic principles
Motion in the atom
The hydrogen nucleus
Net magnetic vector (NMV)
Precession and precessional (Larmor) frequency
The free induction decay (FID) signal
Pulse timing parameters
2 Image weighting and contrast
Relaxation in different tissues
Proton density contrast
3 Spin-echo pulse sequences
Fast or turbo spin-echo (FSE/TSE)
4 Gradient-echo pulse sequences
Variable flip angle
Weighting in gradient-echo pulse sequences
Coherent or rewound gradient-echo
Incoherent or spoiled gradient-echo
Echo planar imaging
5 Spatial encoding
Mechanism of Gradients
Bringing it All Together – Pulse Sequence Timing
Part 1: What is
Part 2: How are Data Acquired and How are Images Created from These Data?
Part 3: Some Important Facts About
Part 4: How Do Pulse Sequences Fill
Part 5: Options that Fill
7 Protocol optimization
Signal-to-Noise Ratio (SNR)
Contrast-to-Noise Ratio (CNR)
Protocol Development and Modification
Chemical Shift Artifact
Out-of-Phase Signal Cancellation
Magnetic Susceptibility Artifact
Flow-Dependent (Non-Contrast-Enhanced) Angiography
Patient Transport System
Computer System and Graphical User Interface
10 MRI safety
Introduction (And Disclaimer)
Definitions Used in MRI Safety
The Spatially Varying Static Field
Electromagnetic (Radiofrequency) Fields
Time-Varying Gradient Magnetic Fields
End User License Agreement
Figure 1.1 The atom.
Figure 1.2 The magnetic moment of the hydrogen nucleus.
Figure 1.3 Alignment – classical theory.
Figure 1.4 Alignment – quantum theory.
Figure 1.5 The net magnetic vector.
Figure 1.6 Precession.
Figure 1.7 Precession of the spin-up and spin-down populations.
Figure 1.8 The electromagnetic spectrum.
Figure 1.9 In phase (coherent) and out of phase (incoherent).
Figure 1.10 Energy transfer during excitation.
Figure 1.11 Phase and frequency (the watch analogy).
Figure 1.12 Generation of the signal.
Figure 1.13 Longitudinal and transverse magnetization.
Figure 1.14 A basic pulse sequence.
Figure 2.1 The T1 recovery curve.
Figure 2.2 T2* decay and field inhomogeneities.
Figure 2.3 Dephasing and free induction decay.
Figure 2.4 The T2 decay curve.
Figure 2.5 The magnitude of transverse magnetization vs amplitude of signal.
Figure 2.6 T1 recovery in fat.
Figure 2.7 T1 recovery in water.
Figure 2.8 T2 decay in fat.
Figure 2.9 T2 decay in water.
Figure 2.10 T1 contrast generation.
Figure 2.11 Saturation with a short TR.
Figure 2.12 No Saturation with a long TR.
Figure 2.13 T2 contrast generation.
Figure 2.14 The difference in T1 recovery between fat and water.
Figure 2.15 Coronal T1-weighted image of the knee.
Figure 2.16 The difference in T2 decay between fat and water.
Figure 2.17 Axial T2-weighted image of the wrist.
Figure 2.18 Sagittal proton density weighted image of the ankle.
Figure 2.19 T1 weighting and the heat analogy.
Figure 2.20 T2 weighting and the heat analogy.
Figure 2.21 Proton density weighting and the heat analogy.
Figure 2.22 Midline sagittal T1-weighted image of the brain.
Figure 2.23 Free and restricted diffusion in water.
Figure 2.24 BOLD images of the brain. Functional areas are shown in red.
Figure 3.1 T2* dephasing.
Figure 3.2 180° RF rephasing.
Figure 3.3 A basic spin-echo pulse sequence.
Figure 3.4 The Larmor Grand Prix.
Figure 3.5 Tau.
Figure 3.6 Spin-echo with one echo.
Figure 3.7 Spin-echo with two echoes.
Figure 3.8 Spatial encoding in conventional spin-echo.
Figure 3.9 The echo train.
Figure 3.10 Phase-encoding gradient slopes.
Figure 3.11 Phase encoding vs signal amplitude.
-Space filling and phase reordering.
Figure 3.13 Sagittal TSE T2-weighted image of a female pelvis. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 3.14 Axial T1-weighted TSE image of a male pelvis. Source: Westbrook 2015 . Reproduced with permission of John Wiley & Sons.
Figure 3.15 Coronal TSE PD-weighted image of a knee. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 3.16 The DRIVE pulse sequence.
Figure 3.17 Axial DRIVE image through the right internal auditory meatus. Note high signal intensity in the CSF.
Figure 3.18 The 180° inverting pulse in an IR pulse sequence.
Figure 3.19 The inversion recovery sequence.
Figure 3.20 T1 weighting in inversion recovery.
Figure 3.21 PD weighting in inversion recovery.
Figure 3.22 Axial T1-weighted inversion recovery sequence of the brain. A T1 of 700 ms was used.
Figure 3.23 The STIR sequence.
Figure 3.24 Coronal STIR of the knee. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 3.25 Axial T2-weighted FLAIR of the brain. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 3.26 Coronal IR-TSE T2-weighted image with a TI selected to null white matter. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 4.1 How gradients dephase.
Figure 4.2 How gradients rephase.
Figure 4.3 A basic gradient-echo sequence showing how a bipolar application of the frequency- encoding gradient produces a gradient-echo.
Figure 4.4 T1 contrast in gradient-echo.
Figure 4.5 T2* contrast in gradient-echo.
Figure 4.6 T1 weighting in gradient-echo and the heat analogy.
Figure 4.7 T2* weighting in gradient-echo and the heat analogy.
Figure 4.8 PD weighting in gradient-echo and the heat analogy.
Figure 4.9 The steady state.
Figure 4.10 Ernst angle graphs in the brain using a TR of 30 ms.
Figure 4.11 Axial steady state image.
Figure 4.12 Echo formation in steady state I.
Figure 4.13 Echo formation in steady state II.
Figure 4.14 The coherent gradient-echo sequence.
Figure 4.15 Axial coherent gradient-echo sequence of the abdomen. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 4.16 Sagittal coherent gradient-echo sequence of the knee. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 4.17 RF spoiling in the incoherent gradient-echo sequence.
Figure 4.18 Sagittal incoherent gradient-echo sequence of the ankle. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 4.19 Coronal incoherent gradient-echo after contrast enhancement. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 4.20 The reverse-echo sequence.
Figure 4.21 Axial reverse-echo gradient-echo in the brain.
Figure 4.22 Perfusion imaging using an echo shifting sequence. Source: Westbrook 2015 . Reproduced with permission of John Wiley & Sons.
Echo formation in coherent gradient-echo.
Echo formation in incoherent gradient-echo.
Echo formation in reverse-echo gradient-echo.
Figure 4.26 Balanced gradient system in balanced gradient-echo.
Figure 4.27 Maintenance of the steady state in balanced gradient-echo.
Figure 4.28 Axial balanced gradient-echo image of the lumbar spine.
Figure 4.29 GE-EPI sequence.
Figure 4.30 SE-EPI sequence.
Figure 4.31 GRASE sequence.
Figure 4.32 Axial SE-EPI of the abdomen. Source: Westbrook 2015 . Reproduced with permission of John Wiley & Sons.
Figure 5.1 How gradients change field strength and precessional frequency.
Figure 5.2 Three-terminal gradient coil.
Figure 5.3 (a) Steep and (b) shallow gradient slopes.
Figure 5.4 Gradient axes in a typical superconducting system.
Figure 5.5 Slice-selection and the tuning fork analogy.
Figure 5.6 X, Y, and Z as slice selectors.
Figure 5.7 Timing of the slice-select gradient in a spin-echo pulse sequence.
Figure 5.8 Transmit bandwidth, gradient slope, and slice thickness.
Figure 5.9 Frequency encoding and the keyboard analogy.
Figure 5.10 Timing of the frequency-encoding gradient in a spin-echo pulse sequence.
Figure 5.11 Phase encoding.
Figure 5.12 Timing of the phase-encoding gradient in a spin-echo pulse sequence.
Figure 5.13 (a) Steep and (b) shallow phase-encoding gradients.
Figure 5.14 Gradient timing in a spin-echo pulse sequence.
Figure 5.15 How a waveform is produced from a change of phase over distance.
Figure 5.16 Spatial frequency vs amplitude of phase-encoding gradient.
Figure 6.1 The axes of
-space – the chest of drawers.
-Space – phase matrix and the number of drawers.
-Space – labeling.
-Space filling in a spin-echo sequence.
Figure 6.6 Waveforms from three different magnet moments precessing at three different frequencies and their amplitude modulation.
Figure 6.7 Data points in
-space. The number in each column is the phase matrix. The number in each row is the frequency matrix.
Figure 6.8 The Nyquist theorem.
Figure 6.9 Sampling time (acquisition window) and the TE.
Figure 6.10 FFT.
Figure 6.11 Signals in
-space from one voxel.
Figure 6.12 Generation of pseudo-frequency from one voxel.
-Space symmetry – phase.
-Space symmetry – frequency.
Figure 6.15 Phase gradient amplitude vs signal amplitude.
-Space – signal and resolution.
-Space using all data.
-Space using resolution data only.
-Space using signal data only.
Figure 6.20 TR vs slice number.
Figure 6.21 How gradients traverse
-space in a gradient-echo sequence.
Figure 6.22 Partial Fourier.
Figure 6.23 Partial echo.
Figure 6.24 Parallel imaging.
-Space filling in EPI.
Figure 6.26 Spiral
Figure 6.27 Data acquisition methods.
Figure 7.1 Coil position vs SNR.
Figure 7.2 (a) TR 700 ms, (b) TR 500 ms, (c) TR 300 ms, (d) TR 140 ms.
Figure 7.3 Changing TR at 3 T.
Figure 7.4 Flip angle vs SNR.
Figure 7.5 Axial gradient-echo image of the brain using a flip angle of 10° at 3 T.
Figure 7.6 Axial gradient-echo image of the brain using a flip angle of 90° at 3 T.
Figure 7.7 SNR vs TE.
Figure 7.8 (a) TE 11 ms, (b) TE 20 ms, (c) TE 40 ms, (d) TE 80 ms.
Figure 7.9 Changing TE at 3 T.
Figure 7.10 SNR vs NSA.
Figure 7.11 Sagittal brain using 1 NSA.
Figure 7.12 Sagittal brain using 4 NSA.
Figure 7.13 SNR vs receive bandwidth.
Figure 7.14 The voxel. The large blue square is the FOV.
Figure 7.15 Voxel volume and SNR (spin numbers are not representative).
Figure 7.16 SNR vs slice thickness.
Figure 7.17 Sagittal brain using 128 phase matrix.
Figure 7.18 Sagittal brain using 256 phase matrix.
Figure 7.19 SNR vs FOV.
Figure 7.20 Sagittal brain using a square FOV of 240 mm.
Figure 7.21 Sagittal brain using a square FOV of 120 mm.
Figure 7.22 A heavily T2-weighted image of the buttock. A very long TE was used in this image. The CNR is optimized showing pathology clearly.
Figure 7.23 Fat saturation.
Figure 7.24 Sagittal ankle with fat saturation. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 7.25 Water saturation.
Figure 7.26 Sagittal T2-weighted image of the pelvis without fat saturation.
Figure 7.27 Sagittal T2-weighted image of the pelvis with fat saturation.
Figure 7.28 Square and rectangular FOV and the chest of drawers.
Figure 8.1 Axial image through a breathing abdomen showing phase mismapping.
Figure 8.2 One of the causes of phase mismapping.
Figure 8.3 Axial T1-weighted image of the chest. Phase is anterior to posterior. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 8.4 Axial T1-weighted image of the chest. Phase is right to left. Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 8.5 Placement of respiratory compensation and cardiac gating leads.
Figure 8.6 Respiratory compensation and
Figure 8.7 Image showing respiratory motion.
Figure 8.8 Image without respiratory motion.
Figure 8.9 (a) Respiratory triggering and (b) gating.
Figure 8.10 Sagittal image of the brain showing aliasing or wrap around.
Figure 8.11 Aliasing and undersampling.
Figure 8.12 Phase wrap.
Figure 8.13 Antialiasing along the frequency axis.
Figure 8.14 Antialiasing along the phase axis.
Figure 8.15 With wrap.
Figure 8.16 Without wrap.
. In this image, the signal from the fatty bone marrow in the talus has been shifted inferiorly in the frequency-encoding direction. This is more apparent on the 16 KHz image where the signal void mimics a thickening of cortical bone.
Figure 8.18 Chemical shift and pixel shift.
Figure 8.19 Out-of-phase signal cancellation show as a black line around the abdominal organs border at the boundaries between fat and muscle.
Figure 8.20 Out-of-phase signal cancellation and the watch analogy.
Figure 8.21 The periodicity of fat and water.
Figure 8.22 Axial gradient-echo images when the magnetic moments of fat and water nuclei are out of phase (below) and in phase (above). Source: Westbrook 2014 . Reproduced with permission of John Wiley & Sons.
Figure 8.23 Magnetic susceptibility from a dental implant causing massive distortion of the image.
Figure 8.24 Sagittal gradient-echo images of the knee with pins in the tibia. Magnetic susceptibility has produced a large distortion of the image.
Figure 8.25 Sagittal spin-echo images of the same patient as shown in Figure 8.24. The artifact is reduced.
Figure 8.26 Truncation artifact.
Figure 8.27 Contrast changes between slices as a result of cross-excitation.
Figure 8.28 Cross-excitation.
. In both examples shown, there is a line of high signal running perpendicular to the frequency-encoding direction. This represents external interference at a discrete frequency.
Figure 8.30 Various appearances of multiple noise spikes.
Figure 8.31 Moiré artifact seen as zebra lines on the edge of the FOV.
Figure 8.32 Magic angle artifact shown with a high signal intensity at the lower border of the patellar tendon.
Figure 8.33 Inflow and flow-related enhancement.
Figure 8.34 Entry-slice phenomenon and direction of flow.
Figure 8.35 Entry-slice phenomenon: slice 1 (most inferior).
Figure 8.36 Entry-slice phenomenon: slice 2 (middle inferior).
Figure 8.37 Entry-slice phenomenon: slice 3 (middle superior).
Figure 8.38 Entry-slice phenomenon: slice 4 (most superior).
Figure 8.39 Time-of-flight phenomenon.
Figure 8.40 Time of flight vs TE.
Figure 8.41 The different types of flow.
Figure 8.42 Intravoxel dephasing.
Figure 8.43 Gradient moment rephasing (nulling).
Figure 8.44 Without GMN.
Figure 8.45 With GMN.
Figure 8.46 Spatial presaturation.
Figure 8.47 Without presaturation.
Figure 8.48 With presaturation.
Figure 8.49 Axial 3D inflow MRA images of the brain to evaluate vasculature in the circle of Willis. These images were acquired at 3 T (a) and 1.5 T (b). Note the improvement in vascular contrast due to greater SNR and CNR in the 3 T image.
. Tortuous vessels or vessels in which the flow is along the plane of the slice may not be demonstrated on a MIP. The spins become saturated by numerous excitation pulses.
MIP reformatting. The maximum intensity projection ascertains the maximum intensity in each row or column of pixels and assigns this value to a pixel in a projected plane. In this diagram, there are two such planes representing an anterior and lateral projection of the data.
Figure 8.52 Image reconstructions from MOTSA acquisition (multiple overlapping thin slabs). Signal saturation has occurred at the edge of a slab resulting in an apparent (and highly suspicious) bilateral “stenosis.”
Figure 8.53 8.53 Spatial presaturation to produce black blood. Note that the magnetic moments of nuclei in the vessel (at the top of the illustration) are aligned with the magnetic field (
) along the
-axis. As the blood within the vessel flows down into the saturation volume, they receive a 90° RF pulse and their vectors enter the transverse plane. As blood continues to flow down into the slice, nuclei receive another 90° RF pulse and are aligned 180° from their original position at the top of the diagram. At this point (and with no time to recover), nuclei with the blood are saturated and are dark on the image.
Figure 8.54 8.54 Black-blood imaging. Normal flow may return a similar signal intensity to the vessel lumen. In the event of an arterial dissection, stationary blood trapped in the intima may also exhibit a similar intensity. By saturating inflowing blood, the signal is removed resulting in a more accurate and conspicuous representation of vessel patency.
Figure 8.55 PCA gradient.
Figure 8.56 VENC.
Figure 8.57 VENC aliasing.
Figure 8.58 8.58 EKG-triggered subtraction imaging. This technique exploits the T2 contrast difference between blood (a fluid) and the background tissue. Data acquisition is synchronized to the EKG permitting the visualization of arteries or veins. The mechanism is related to the speed of flow.
Figure 9.1 Closed-bore MRI scanner in axial cross-section revealing the principal components to be arranged in concentric circles, most of them being cylindrical electromagnets.
Figure 9.2 Effect of a diamagnetic substance on a homogenous magnetic field.
Figure 9.3 Effect of a paramagnetic substance on a homogenous magnetic field.
Figure 9.4 Ferromagnetic substance in a homogenous magnetic field.
Figure 9.5 Differences in solenoid configuration in (a) closed-bore and (b) open MRI scanners.
Figure 9.6 Typical design of a permanent-magnet open scanner. The flux lines of the static field run vertically in this type of scanner.
Figure 9.7 Right-hand grip rule.
Figure 9.8 Construction of an MRI cryostat.
Figure 9.9 Bobbin used to create a segmented solenoid electromagnet.
Figure 9.10 Parallel circuit and persistent switch used for ramping a magnet.
Figure 9.11 Passive shielding causes the lines of flux to pass through steel cladding in preference to air.
Figure 9.12 Passive shim system.
Figure 9.13 Mechanism of a gradient coil.
Figure 9.14 How gradients change field strength.
Figure 9.15 Characteristics of a magnetic field gradient.
Figure 9.16 RF transmit and receive chain.
Figure 9.17 The RF transmit coil. This is a birdcage resonator coil consisting of two end rings linked by a symmetrical array of straight conductors.
Figure 9.18 Schematic diagram showing surface coil configurations.
Figure 9.19 Phased array spine coil.
Figure 10.1 MRI safety zones as recommended by the ACS Guidance Document on MR-Safe Practices 2013.
Figure 10.2 Device-labeling icons developed by the American Society for Testing and Materials International and recognized by the FDA.
Figure 10.3 Static magnetic field gradient of an actively shielded closed-bore MRI scanner. The values shown will vary with scanner model.
Figure 10.4 Ferromagnetic wheelchair taken into an actively shielded magnet room. No attraction was appreciated until the chair was moved close to the end of the bore.
Figure 10.5 An implanted device or clip may experience torque. This is a turning force that causes the long axis of a device to become aligned with the static field.
Figure 10.6 Anatomy may form “biological circuits” through which an induced current may flow. Contact areas such as the hands or heels may concentrate the current flow through a small area causing burns.
Figure 10.7 Dipole heating in a conductive wire such as a pacemaker lead.
Table of Contents
The MRI in Practice brand continues to grow from strength to strength. The fourth edition of MRI in Practice is an international best-seller and is translated into several languages. At the time of writing, the accompanying MRI in Practice course is 26 years old. We have delivered the course to more than 10 000 people in over 20 countries and have a large and growing MRI in Practice online community. Our readers and course delegates include a variety of professionals such as radiographers, technologists, radiologists, radiotherapists, veterinary practitioners, nuclear medicine technologists, radiography students, postgraduate students, medical students, physicists, and engineers.
The unique selling point of MRI in Practice has always been its user-friendly approach to physics. Difficult concepts are explained as simply as possible and supported by clear diagrams, images, and animations. Clinical practitioners are not usually interested in pages of math and just want to know how it essentially “all works.” We believe that MRI in Practice is so popular because it speaks your language without being oversimplistic.
This fifth edition has had a significant overhaul and specifically plays to the strengths of the MRI in Practice brand. We have created a synergy between the book and the course so that they are best able to support your learning. We purposefully focus on physics in this edition and on essential concepts. It is important to get the fundamentals right, as they underpin more specialist areas of practice. There are completely new chapters on MRI equipment and safety, and substantially revised and expanded chapters on gradient-echo pulse sequences, k-space, artifacts, and angiography. The very popular learning tips and analogies from previous editions are expanded and revised. There is also a new glossary, lots of new diagrams and images, and suggestions for further reading for those who wish to delve deeper into physics. The accompanying website includes new questions and additional animations. We also include some equations in this edition, but don’t worry: they are there only for those who like equations, and we explain what they mean in a user-friendly style.
However, probably the most significant change in this edition is the inclusion of scan tips. Throughout the book, your attention is drawn to how theory applies to practice. Scan tips are specifically used to alert you to what is going on “behind the scenes” when you select a parameter in the scan protocol. We hope this helps you make the connection between theory and practice. Physics in isolation is of little value to the clinical practitioner. What is important is how this knowledge is applied. We stand by the MRI in Practice philosophy that physics does not have to be difficult, and we hope that our readers, old and new, find these changes helpful. Richard Feynman, who is considered one of the finest physics teachers of all time, was renowned for his ability to transfer his deep understanding of physics to the page with clarity and a minimum of fuss. He believed that it is unnecessary to make physics more complicated than it need be. Our aspiration is that the fifth edition of MRI in Practice emulates his way of thinking.
We hope that the many fans of MRI in Practice around the world continue to enjoy and learn from it. A big thank you for your continued support and happy reading!
Many thanks to all my loved ones for their continued support, especially Maggie Barbieri (my mother, whose brain scans feature many times in all the editions of this book and in the MRI in Practice course for the last 26 years. She must have the most viewed brain in the world!), Francesca Bellavista, Amabel Grant, Adam, Ben and Maddie Westbrook.
I’d like to thank my family Dannie, Joey, and Harry for bringing coffee, biscuits, and occasionally gin and tonic. I would also like to take the opportunity to acknowledge the work of a great MRI pioneer, Prof. Sir Peter Mansfield, who died this year. Prof. Mansfield’s team created the first human NMR image in 1976, and he kindly shared all of his most important research papers with me when I first started writing about this amazing field.
Conventional spin-echo (SE)
Turbo spin-echo (TSE)
Single-shot TSE (SS-TSE)
TSE (with restoration pulse)
driven equilibrium FSE
T2 Puls FSE
Inversion recovery (IR)
Fast inversion recovery
Short tau IR (STIR)
Fluid-attenuated IR (FLAIR)
turbo dark fluid
Echo-planar imaging (EPI)
Double-echo steady state
Balanced dual excitation
phase balanced SARGE
fast GRE, fast SPGR
Repetition time (TR)
Time to echo (TE)
Time from inversion (TI)
Number of echoes (in TSE)
Field of view (FOV)
Data acquisition parameters
receive bandwidth (KHz)
fat water shift (pixel)
Parallel imaging (image based)
Parallel imaging (
Artifact reduction techniques
Gradient moment rephasing
Moving sat pulse
Sequential pre SAT
frequency wrap suppression
no phase wrap
phase wrap suppression
Volume TSE variable flip angle
keyhole (4d Trak)
Noncontrast MRA gradient-echo
NATIVE – true FISP
inhance inflow IR
Noncontrast MRA spin- echo
High-resolution breast imaging
Diffusion tensor imaging
diffusion tensor imaging
Body diffusion imaging
spin quantum number
number of spins in the high-energy population (Boltzmann)
number of spins in the low-energy population (Boltzmann)
energy difference between high- and low-energy populations (Boltzmann)
temperature of the tissue
precessional or Larmor frequency
external magnetic field strength
energy of a photon
precessional frequency of
magnetic field associated with the RF excitation pulse
duration of the RF excitation pulse
number of turns in a coil
changing magnetic flux in a single loop
amount of longitudinal magnetization at time
full longitudinal magnetization
amount of transverse magnetization at time
full transverse magnetization
signal intensity in a tissue
variation in magnetic field
time between two gradient pulses
echo spacing in turbo spin-echo (TSE)
time from inversion (TI)
TE set at the console
magnetic field strength at a point along the gradient
digital sampling frequency
number of slice locations
max amplitude of the phase encoding gradient
incremental step between each
amplitude of the frequency encoding gradient
standard deviation of background signal or noise
separation between ghosts due to motion p
period of motion of something moving in the patient
density of blood
velocity of flow
diameter of a vessel
viscosity of blood
chemical shift frequency difference between fat and water
chemical shift (3.5 ppm or 3.5 × 10
charge of a particle
Lorentz force (total emf on a charged particle)
electric field vector
magnetic field vector
This book is accompanied by a companion website:
The website includes:
Brand new 3D animations of more complex concepts from the book
100 short-answer questions to aid learning and understanding.
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