Silke Hecht
Basic MRI physics135, 142, 235, 278, 351, 354, 385
Magnetic resonance imaging (MRI) is the imaging modality of choice for the diagnosis of most neurologic diseases in human and veterinary patients. While an in-depth review of MR physics is beyond the scope of this chapter, an overview of basic princi-ples and sequences will be provided to facilitate understanding of clinical applications and interpretation.
In general, any element with either an odd number of protons or an odd number of neutrons has a nuclear magnetic dipole moment and may therefore be suitable for MR imaging or spec-troscopy. Hydrogen is the optimal element for MR imaging as (1) it is the most common element in the body, (2) its nucleus consists of a single proton and has the strongest magnetic dipole moment of any element suitable for MR imaging, and (3) most pathologic processes affecting the central nervous system (CNS) result in alteration of content, distribution, and ambient envi-ronment of hydrogen protons facilitating differentiation of dis-eased from normal tissue. This overview will focus on MR imag-ing of hydrogen protons.
Hydrogen protons in tissues are not static but spin around their axes, generating their own micromagnetic environments.
In the absence of an external magnetic field the magnetic moments of the spinning protons are randomly oriented. How-ever, when brought into a strong external magnetic field (i.e. an MR scanner), they rearrange under its influence. The magnetic field strength is denoted by the unit “tesla” (T). The strength of clinically used MR scanners ranges from approximately 0.2T (low-field) to 3T (high-field); higher-strength scanners (7–21T) are at this point limited to research institutions. The alignment of
individual protons may be parallel or antiparallel with the exter-nal magnetic field. A slight majority of protons will align with the magnetic field, generating a magnetic vector (“net magneti-zation vector”) which is utilized during MR imaging. The main magnetic field is denoted by B0, the tissue magnetization vector by M0. As long as M0is parallel with the much stronger B0it cannot be easily separated out and cannot be used for imaging.
The goal of MR imaging is to manipulate tissue magnetization in a way that it can be distinguished from the external magnetic environment.
In addition to a spinning motion around their individual axes, hydrogen protons wobble (or precess) under the influence of B0, similar to a spinning top wobbling under the influence of gravity.
The precession frequency of the spinning protons is dependent on the strength of the external magnetic field and is described by the Larmor equation:
ω0= γB0 where:
ω0=Larmor or resonant frequency
γ =gyromagnetic ratio specific for each MR active nucleus (42.56 MHz/T for hydrogen)
B0=external magnetic field strength.
In order to manipulate tissue magnetization so it can be sepa-rated from the main magnetic field, radiofrequency (RF) pulses are applied. Once the nuclei are exposed to an RF pulse exactly matching their precessional frequency they gain energy and start resonating. As a result of the energy gain some protons change their alignment with the magnetic field, causing the magnetic net vector to move away from B0, or “flip.” The most common flip angle of the tissue magnetization vector is 90◦used in spin echo sequences (see below). Hydrogen protons flipped into this
“transverse plane”—which is in perpendicular orientation to the main magnetic field (the “longitudinal plane”)—continue to pre-cess. According to Faraday’s law, any change in the magnetic environment of a coil of wire will cause an electric signal. Strate-gic placement of receiver coils in the MR unit allows detec-tion and measurement of magnetizadetec-tion in the transverse plane, which is the basis of image formation in MRI (Fig. 6.1).
Practical Guide to Canine and Feline Neurology, Third Edition. Edited by Curtis W. Dewey and Ronaldo C. da Costa.
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(a)
(b)
Figure 6.1 Schematic illustration of the alignment of the net vector (tissue magnetization) before (A) and after (B) application of a 90◦RF pulse. (A) The net vector is aligned parallel to the main magnetic field (B0) and cannot be measured. (B) After application of the 90◦pulse the net vector is oriented perpendicular to the main magnetic field, and a current (signal) is induced in the receiver coil. (The Ohio State University. Reproduced with permission.)
After the RF pulse is switched off, the signal induced in the receiver drops off rapidly due to two concurrent processes:
T1-relaxation (spin-lattice relaxation): Excited hydrogen pro-tons return to lower energy states, and the magnetic net vec-tor realigns with the main magnetic field (Fig. 6.2).
T2-relaxation (spin-spin relaxation): The unified front of hydrogen protons quickly loses coherence, resulting in a
signal drop-off. While true T2-relaxation is solely the result of micromagnetic inhomogeneities (spin–spin interactions, i.e. interference of one spinning proton’s micromagnetic field with its neighbors), extrinsic magnetic field inhomogeneities (magnet imperfections, disruption of the magnetic field by paramagnetic or ferromagnetic substances, etc.) contribute to an even more rapid loss of phase coherence. This process is called T2*-relaxation (Fig. 6.3).
Three important tissue parameters for MR imaging include T1- and T2-relaxation times and proton density (PD), i.e. the actual amount of hydrogen protons in a certain tissue volume.
The goal of MR imaging is to translate these inherent tissue ferences into image contrast, which is accomplished by using dif-ferent MR sequences.
MR sequences
Basic spin echo (SE) sequences115, 179, 274, 279, 285–287, 293, 385
These include T1-W, T2-W and proton density (PD) weight-ing, which are the most basic but also the most commonly used MRI sequences. Each SE sequence starts with a 90◦RF pulse fol-lowed by a 180◦pulse applied exactly halfway between the initial 90◦pulse and the generation of the signal (echo) (Fig. 6.4). The 180◦pulse is applied to cancel out external magnetic field inho-mogeneities. It essentially reverses any effects an external dis-turbance (e.g. a nearby flowing vessel or paramagnetic methe-moglobin in a hematoma) will have on proton alignment and resultant tissue signal. The small interactions occurring between individual protons and contributing to signal loss in the trans-verse plane cannot be retrans-versed, resulting in true T2 relaxation contributing to image contrast. As the MR signal generated dur-ing a sdur-ingle episode of proton excitation is too small to create an image, the process is repeated many times until enough data have been collected. The time between the 90◦ pulse and the echo is called “time of echo” (TE) and the time between succes-sive 90◦pulses is called “time of repetition” (TR). The length of TR and TE determines weighting of an SE sequence (Table 6.1).
T1-weighting (T1-W): A short TR is chosen to maximize the differences in T1 relaxation between tissues (Fig. 6.5).
This is combined with a short TE to minimize T2 effects.
(a) (b) (c) (d)
Figure 6.2 Schematic illustration of T1 relaxation. (A) Situation before the 90◦pulse. The magnetic vector is aligned with the main magnetic field. (B) After the 90◦pulse the vector is in the transverse plane. (C, D) The protons realign with the magnetic field. (The Ohio State University. Reproduced with permission.)
(a) (b) (c) (d)
Figure 6.3 Schematic illustration of T2/T2* relaxation. (A) Situation before the 90◦pulse. The magnetic vector is aligned with the main magnetic field. (B) After the 90◦pulse the vector is in the transverse plane, the protons form a unified front, and the MR signal is strongest. (C, D) The precessing protons lose coherence, and the strength of the MR signal decreases. This occurs due to interference of individual protons with each other (T2 relaxation) as well as influence of external field inhomogeneities (T2* relaxation). (The Ohio State University. Reproduced with permission.)
Figure 6.4 Schematic outline of a spin echo sequence. (The Ohio State University. Reproduced with permission.)
TR=time of repetition; TE=time of echo; RF=radiofrequency pulse;
SEG/PEG/FEG=slice/phase/frequency encoding gradient (applied for spatial encoding of the MR signal); DAQ=Data acquisition.
Table 6.1 Influence of acquisition parameters on image contrast in a spin echo sequence.36, 351, 385
TR TE Weighting
Short (300–700ms) Short (5–30ms) T1-W
Long (2000–4000ms) Long (60–150ms) T2-W
Long (>2000–4000ms) Short (5–30ms) PD-W
Fat has a short T1 relaxation time and is hyperintense, while fluid has a long T1 relaxation time and appears hypointense. Soft tissues have somewhat variable, inter-mediate T1 relaxation times and are medium in intensity.
After uptake of administered paramagnetic contrast agents, physiologically contrast-enhancing tissues (e.g. pituitary gland) and contrast-enhancing pathologic lesions (e.g.
certain brain tumors) are hyperintense (Fig. 6.6).
T2-weighting (T2-W):A long TE is chosen to maximize differ-ences in T2 relaxation between tissues, combined with a long TR to minimize T1 relaxation effects. A long TE ensures that
(a) (b) (c) (d)
Figure 6.5 Schematic demonstration of a T1-W sequence. (A) Initially, the net vector of all hydrogen protons (fat, soft tissue, fluid) is aligned with the main magnetic field. (B) Immediately after application of a 90◦RF pulse all tissues are in the transverse plane but are quickly separated due to differences in realignment with the main magnetic field (T1 relaxation). Fat has a very short T1 relaxation time, fluid has the longest relaxation time, and soft tissues are intermediate. (C) The second 90◦RF pulse is applied when fat protons have realigned with the main magnetic field. (D) Fat is now alone in the transverse plane and gives the strongest MR signal, i.e. it appears bright on the resultant image. Fluid has the least transverse magnetization and appears dark; soft tissue is intermediate in its intensity. (The Ohio State University. Reproduced with permission.)
(b) (a)
Figure 6.6 Transverse pre- (A) and post- (B) T1-W images of the brain of a dog with a large plaque-like meningioma of the left frontal, parietal, and temporal lobes. Fat associated with subcutaneous tissue and bone marrow and contrast-enhancing tissues appear hyperintense, fluid is hypointense, and soft tissues are intermediate. Additional mass effect beyond the margins of the contrast-enhancing lesion is indicated by midline shift and compression of the left lateral ventricle. The underlying reason (vasogenic edema) is not clearly identified using this sequence (see Fig. 6.8 and Fig. 6.10).
tissues with a short T2 relaxation time will have completely lost their transverse magnetization and have a low signal at the time of image acquisition, while tissues with longer T2 relaxation times still maintain their transverse magnetization and appear brighter (Fig. 6.7). Fluid has a long T2 relaxation time and therefore is hyperintense on T2-W images. Soft tis-sues have intermediate T2 relaxation times. Fat has a short T2 relaxation time and appears hypointense on conventional T2-W images. However, as conventional T2-W SE sequences have largely been replaced with shorter fast spin echo (FSE)
or turbo spin echo (TSE) sequences in which additional pulses are applied (see below), fat typically appears hyperin-tense on today’s T2-W MRI studies. A T2-W sequence can be considered a “pathology” scan, because abnormal fluid collections and tissues with abnormal increased fluid con-tent (“juicy tissue,” e.g. edema, inflammation, neoplasia) will appear hyperintense compared to normal tissues (Fig. 6.8).
Proton density weighting (PD-W): PD-W is achieved by choosing a long TR in combination with a short TE to min-imize T1 and T2 effects on image contrast. PD-W images
(a) (b) (c) (d)
Figure 6.7 Schematic demonstration of a T2-W sequence. (A) Initially, hydrogen protons in the same tissue are aligned with the main magnetic field and are in phase with each other. (B) Immediately after the 90◦RF pulse hydrogen protons in the same substance (fat, soft tissue, fluid) precess in sync with each other in the transverse plane, and all tissues give a strong and uniform signal. (C) Spinning protons quickly lose their coherence due to interference with each other (T2 relaxation). Fluid has the longest T2 relaxation time; fat the shortest. (D) If waiting a long time until listening for the echo (long TE), protons in fat and soft tissues have lost most of their phase coherence. Fluid maintains the strongest transverse magnetization and appears bright on the resultant image. For SE sequences a 180◦pulse is applied at TE/2 to cancel out external field inhomogeneities. (The Ohio State University. Reproduced with permission.)
Figure 6.8 Transverse T2-W image of the brain of a dog with meningioma (same dog as Fig. 6.6 and Fig. 6.10). Fluid within the ventricular system is strongly hyperintense. The peripheral plaque-like mass is isointense to hyperintense to normal brain parenchyma. Extensive T2 hyperintensity to the white-matter tracts of the left cerebral hemisphere and associated mass effect are consistent with vasogenic brain edema.
are characterized by excellent anatomic detail (see Fig. 6.41A below) and are very useful in orthopedic imaging. They also provide a good contrast between gray and white matter, and although their value in the neuroimaging of small animals is limited, anecdotal evidence suggests they may be helpful in the evaluation of patients with degenerative brain disease.
Modified spin echo (SE) sequences23, 50, 83, 94, 95, 131, 179, 217, 269, 293, 309
These sequences are based on conventional SE principles, but additional pulses are applied to selectively suppress signal from certain tissues (inversion recovery sequences) or to accelerate data acquisition (FSE or TSE techniques and single-shot tech-niques).
Inversion recovery sequences: These are characterized by an initial 180◦pulse (inversion pulse). Dependent on the time elapsed between this 180◦ pulse and initiation of the regu-lar SE sequence (time of inversion; TI), they result in selec-tive suppression of fluid (fluid attenuated inversion recovery;
FLAIR) or fat (short tau inversion recovery; STIR).
FLAIR:A long TI prior to initiation of a SE sequence allows selective suppression of fluid (Fig. 6.9). Although FLAIR images are most commonly used for brain imaging, they may occasionally be helpful in the evaluation of spinal lesions.
T2-W FLAIR images are useful in conjunction with regu-lar T2-W images in characterizing T2 hyperintense lesions.
Using FLAIR, pure fluid (cerebrospinal fluid and fluid in cys-tic lesions) is suppressed and becomes hypointense, while solid lesions remain hyperintense (Fig. 6.10). Additionally, this sequence increases conspicuity of small lesions border-ing a fluid-filled ventricle or the subarachnoid space. Finally, FLAIR is helpful in differentiating true T2 hyperintense parenchymal lesions from pseudolesions created by inclusion of fluid-filled structures and brain parenchyma within the same slice thickness (volume averaging). Without modifica-tion of acquisimodifica-tion parameters, FLAIR is unable to suppress signal from fluids containing high-protein cell components or blood by-products, a potential pitfall when interpreting
(a) (b) (c) (d)
Figure 6.9 Schematic demonstration of a FLAIR sequence. (A) Initially, the net vector of all hydrogen protons (fat, soft tissue, fluid) is aligned with the main magnetic field. (B) Immediately after application of the initial 180◦RF pulse (“inversion pulse”), the vectors of different tissues are flipped out of plane together but are quickly separated due to differences in T1 relaxation. Fat has a very short relaxation time, fluid has the longest relaxation time, and soft tissues are intermediate. (C) The 90◦RF pulse is applied when fat has almost realigned with the main magnetic field and fluid has reached the transverse plane. (D) Fluid is now furthest away from the transverse plane and gives off no signal, i.e. it is effectively nulled. (The Ohio State University.
Reproduced with permission.)
Figure 6.10 Transverse T2-W FLAIR image of the brain of a dog with meningioma (same dog as Fig. 6.6 and Fig. 6.8). Pure fluid within the ventricular system is attenuated, while hyperintensity to the white-matter tracts of the left cerebral hemisphere consistent with vasogenic edema persists. The plaque-like peripheral mass is isointense to adjacent tissues and not clearly seen.
images. Postcontrast T1-W FLAIR images are very sensitive in the detection of contrast-enhancing lesions and may be used as an alternative to conventional postcontrast T1-W SE images.
STIR:A short TI prior to initiation of an SE sequence allows selective suppression of fat (Fig. 6.11). This sequence is very
valuable in orthopedic and spinal imaging as it allows dif-ferentiation of pathologic T2 hyperintense lesions within the vertebral canal, vertebrae, and surrounding paraspinal tissues from fat (Fig. 6.12).
Fast (FSE) or turbo (TSE) spin echo techniques:In conven-tional SE imaging one 180◦ pulse is applied during each TR, and one echo (signal) is generated. In FSE and TSE, multiple 180◦ pulses are applied during each TR and mul-tiple echoes are received, resulting in a decrease in scan time without compromising image quality. FSE/TSE techniques have essentially replaced conventional SE sequences in T2-W imaging. One potential disadvantage is strong hyperintensity of fat on T2-W FSE/TSE images as hyperintense epidural fat in the vertebral canal may obscure T2-hyperintense lesions within adjacent spinal cord, and hyperintense fat in bone marrow may obscure or mimic skull or vertebral lesions. This disadvantage can easily be compensated for by adding a fat-saturation technique or comparing T2-W images to other sequences (see Fig. 6.12).
Single-shot techniques: These ultrafast techniques employ a single RF pulse, further decreasing scan time. The resul-tant images are characterized by very strong T2 contrast and are most beneficial in imaging of fluid-filled spaces. “Quick-brain” MR imaging was initially introduced as an alterna-tive technique to CT scanning for assessing children with hydrocephalus. Other indications in humans include macro-cephaly, Chiari malformation, intracranial cysts, screening prior to lumbar puncture, screening for congenital anoma-lies, and trauma. These sequences have gained popularity in veterinary medicine for spinal imaging due to their myelo-graphic effect, which can be used to classify spinal lesions, identify sites of significant intervertebral disc herniation and diagnose spinal subarachnoid cysts (Fig. 6.13).
(a) (b) (c) (d)
Figure 6.11 Schematic demonstration of a STIR sequence. (A) Initially, the net vector of all hydrogen protons (fat, soft tissue, fluid) is aligned with the main magnetic field. (B) Immediately after application of the initial 180◦RF pulse (“inversion pulse”), the vectors of different tissues are flipped out of plane together but are quickly separated due to differences in T1 relaxation. Fat has a very short relaxation time, fluid has the longest relaxation time, and soft tissues are intermediate. (C) The 90◦RF pulse is applied when fat has reached the transverse plane while fluid and soft tissue lag behind. (D) Fat is now furthest away from the transverse plane and gives off no signal, i.e. it is effectively nulled. (The Ohio State University. Reproduced with permission.)
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Figure 6.12 Sagittal T2-W FSE (A) and STIR (B) images of the thoracolumbar spine in a dog following trauma. Traumatic intervertebral disc herniation and mild subluxation at T12/13 along with multifocal disc degenerative changes are evident on both sequences. (A) Multifocal T2 hyperintensities within the vertebral bodies (most obvious in the caudal half of L2; arrowhead) are seen. (B) Most vertebral hyperintensities are attenuated (arrowhead), indicating that they are consistent with fatty degeneration of bone marrow. The cranial half of T13 remains hyperintense, consistent with bone contusion. Additionally, extensive subcutaneous edema dorsal to the lumbar spine (*) and patchy hyperintensities in the epaxial muscles dorsal to T12/13 are accentuated due to suppression of adjacent fat.
(a)
(b)
Figure 6.13 Sagittal T2-W FSE (A) and half-Fourier-acquisition single-shot turbo spin echo HASTE (B) images of a dorsal subarachnoid cyst/diverticulum at T13/L1. The lesion is accentuated on the heavily T2-W (myelographic) single-shot sequence.
Figure 6.14 Schematic outline of a gradient-echo sequence. In contrast to SE techniques only a single RF pulse is applied which is used in conjunction with gradient reversals. (The Ohio State University.
Reproduced with permission.)
TR=time of repetition; TE=time of echo; SSG/PEG/FEG=
slice/phase/frequency encoding gradient (applied for spatial encoding of the MR signal).
Gradient recalled echo (GRE)36, 81, 94, 104, 110, 131, 179, 208, 224, 259, 265, 279, 286, 289, 294, 319, 350, 352, 366, 383–385, 387
While the generation of SE images relies on the application of pairs of RF pulses, GRE sequences utilize only one initial RF pulse in conjunction with gradient field reversals (Fig. 6.14).
GRE sequences use smaller flip angles and shorter TRs than SE sequences (Table 6.2), resulting in shorter scan times. The lack of a 180◦ pulse has important implications for image weight-ing and quality: (1) conventional GRE sequences can be used to acquire T1-W, PD-W and T2*-W images—acquisition of truly T2-W images is not possible—and (2) GRE sequences are prone to susceptibility artifacts as there is no compensation for external field inhomogeneities. A plethora of GRE applications have been developed in recent years, including conventional, steady-state, coherent, incoherent (spoiled), steady-state-free precession, bal-anced, fast, single-shot and echoplanar imaging (EPI) sequences.
A comparison between sequences developed by different ven-dors is difficult as a uniform nomenclature system does not exist.
For example, vendor-specific names/acronyms used for a similar coherent gradient echo sequence are “FISP” (Siemens), “GRASS”
(GE), “FFE” (Philips), “Rephased SARGE” (Hitachi) and “SSFP”
(Toshiba). Rapid development in the field of GRE sequences led to numerous new and advanced applications of MR imaging,
Table 6.2 Influence of acquisition parameters on image contrast in a gradient echo sequence.37, 385
TR TE Flip angle Weighting
Short (<50ms) Short (<5ms) Large (70–90◦) T1-W Long (>100ms) Long (15–25ms) Small (5–20◦) T2*-W Long (>100ms) Short (<5ms) Small (5–30◦) PD-W
(b) (a)
Figure 6.15 Transverse T2-W FSE (A) and T2*-W GRE (B) images of the brain in a dog with a hemorrhagic infarct. (A) An irregular heterogeneous T2 hyperintense mass is associated with the right frontal lobe. Diffuse T2 hyperintensity extending along the adjacent white-matter tracts is consistent with vasogenic edema. (B) The mass is strongly hypointense on T2*-W image, consistent with susceptibility artifact and indicative of hemorrhage.
such as motion-free (breath hold) abdominal imaging, 3D vol-umetric imaging, and 3D MR angiography (MRA). Some GRE sequences more or less routinely used for the imaging of the CNS in small animals at this point include T2*-W GRE, T1-W 3D GRE, 2D/3D volumetric acquisition, and magnetic resonance angiography (MRA).
T2*-W GRE sequence:Gas interfaces, soft-tissue mineraliza-tion, fibrous tissue, and certain blood degradation products (e.g. methemoglobin) cause magnetic field inhomogeneities which appear as a signal void (susceptibility artifact, see below) on T2*-W images. T2*-W is most commonly utilized to identify intracranial or spinal hemorrhage and to differ-entiate it from other lesions (Fig. 6.15; see also Figs 6.38, 6.41B and 6.50). Additional indications may include identi-fication of intracranial mineralization (e.g. in meningiomas) or abnormal gas pockets (e.g. in brain abscesses).
T1-W 3D GRE:These sequences may be beneficial to evalu-ate small structures (e.g. pituitary gland or cranial nerves;
Fig. 6.16) after intravenous administration of contrast medium, as they allow acquisition of thin slices (<1 mm) without interslice gap and permit multiplanar reconstruction of the 3D dataset in additional planes.
2D/3D volumetric acquisition (steady-state-free precession):
These sequences are characterized by high contrast between fluid-filled structures and surrounding tissues and may be beneficial in evaluating small structures, such as the inner ear.
Magnetic resonance angiography (MRA): MRA techniques maximize vascular contrast by enhancing the signal from spins in flowing blood and/or suppressing the signal from surrounding stationary tissues. Although this can be accom-plished without contrast medium administration (digital subtraction MRA, phase contrast MRA, time-of-flight MRA) contrast-enhanced MRA (CE-MRA) is considered a superior
technique due to improved image quality. In humans, evalu-ation of the intracranial circulevalu-ation provides valuable infor-mation in the diagnosis and prognosis of various abnormal-ities such as aneurysms, arterial and venous steno-occlusive diseases, inflammatory arterial diseases, and congenital vas-cular abnormalities. Although intracranial vasvas-cular abnor-malities are infrequently reported in the veterinary literature, MRA might be considered a quick and low-risk procedure to evaluate intracranial vessels in select cases.
Figure 6.16 Postcontrast transverse T1-W GRE (volume interpolated breath hold examination; VIBE) image of the brain in a dog with left-sided facial paralysis (slice thickness 0.9 mm). Contrast enhancement of the left facial nerve is noted (arrow), consistent with facial neuritis.
Functional imaging18, 25, 37, 45, 48, 62, 126, 142, 143, 154, 194, 223, 238, 253, 268, 284, 297, 334, 352, 370, 381, 385, 399
Diffusion-weighted imaging (DWI): Diffusion describes the motion of molecules in tissues. This process is not truly random due to the presence of physiologic boundaries (cell membranes etc.) and is referred to as “apparent diffusion.”
DWI utilizes opposing gradients to produce signal differ-ences based on mobility and direction of water diffusion.
Normal tissues have more water mobility, resulting in greater signal loss, while tissues with less water mobility experience restricted diffusion. In human as well as veterinary medicine DWI is most commonly used in the diagnosis of ischemic stroke. In acute cerebral ischemia, restricted diffusion occurs secondary to failure of the cell membrane ion pump and sub-sequent cytotoxic edema. An acute stroke is characterized by marked hyperintensity on a DWI and hypointensity on a syn-thesized apparent diffusion coefficient (ADC) map derived from two or more DWIs (Fig. 6.17). In humans DWI is also used to differentiate benign from malignant lesions and dis-tinguish neoplasia from edema or infarction. While some initial studies in animals have yielded promising results, the value of DWI in the diagnosis of neurologic disorders other than acute stroke remains to be determined. Diffusion tensor imaging (DTI) is a specialized DWI technique which utilizes strong multidirectional gradients to map white matter tracts (Fig. 6.18). Initial studies proved feasibility of this technique in dogs, which may ultimately aid in the diagnosis of white matter disease and facilitate surgical planning.
Perfusion-weighting imaging (PWI):PWI allows an estimate of blood volume passing through the capillary bed per unit of time. This is most commonly accomplished by trac-ing the passage of a bolus of contrast agent through the cerebral vasculature. Perfusion imaging is often used in
combination with DWI in patients with acute ischemic stroke, where the difference between diffusion and perfusion abnormalities provides a measure of the ischemic penum-bra (area of reversible ischemia that can be salvaged if blood flow is re-established promptly). Other potential applications include assessment of tumor malignancy based on metabolic activity and evaluation of tissue viability of vascular organs.
Functional MR imaging (fMRI):This is a rapidly evolving area in human medicine which allows evaluation of brain activity during certain activities or following stimulation. Any activ-ity/stimulus (e.g. viewing a picture or smelling food) results in activation of a specific area in the brain, which necessi-tates an increase in blood flow. The resultant focal increase in oxyhemoglobin and decrease in deoxyhemoglobin can be detected by means of MRI due to their inherent differ-ence in magnetic susceptibility (blood oxygen level depen-dent, or BOLD, imaging). The result is a map of functional brain areas during a specific activity or after a specific stimulus (Fig. 6.19). Unfortunately, application of this tech-nique in veterinary patients is thus far limited due to the need for immobilization or general anesthesia. However, ini-tial attempts in conditioned awake dogs yielded promising results, and fMRI in animals is likely to gain importance with development of faster sequences and experience.
Magnetic resonance spectroscopy (MRS):MRS is a method to measure tissue chemistry by recording signals from spe-cific metabolites. In vivo MRS is most commonly per-formed using hydrogen protons (1H [proton] spectroscopy).
Although other metabolites can be recorded using this technique, proton spectroscopy has the advantages of a high signal-to-noise ratio and a relatively short examina-tion time as it can be added to convenexamina-tional MR imaging protocols. Applications in neuroimaging include monitor-ing of biochemical changes occurrmonitor-ing in tumors, metabolic
(a) (b) (c)
Figure 6.17 Ischemic cerebellar infarct in a dog. (A) The transverse T2-W image demonstrates a sharply marginated wedge-shaped hyperintense lesion associated with the left cerebellar hemisphere. (B, C) The lesion remains hyperintense on the diffusion-weighted image (B) and is hypointense on the ADC map (C), consistent with restricted diffusion and ischemic stroke.
Figure 6.18Diffusion tensor imaging of the corticospinal tract in a dog (sagittal view).
(Jacqmotet al., 2013. Reproduced with permission from Wiley.)143
disorders, inflammatory, and neurodegenerative diseases.
With increasing availability of higher-strength magnets, MRS is gaining popularity in clinical veterinary medicine.
Technical modifications34, 56, 64, 65, 70, 99, 370, 373
Spatial presaturation:Spatial presaturation pulses are used to suppress undesired signals from anatomic areas within the imaging field of view. These pulses are not commonly applied in brain imaging but may be helpful in spinal imaging where they can be used to suppress signal from neighboring vessels or peristaltic bowel, thus minimizing motion artifacts.
Figure 6.19 Functional MRI (fMRI) demonstrating increased activity in the area of the right caudate nucleus (CD) in response to a hand signal (image summed from data acquisition in two awake dogs). (From Berns et al., 2012. Reprinted with permission.)25
Fat saturation:Unlike STIR, which is an entirely separate MR sequence, selective fat saturation pulses can be applied to any sequence to suppress signal from fat without affecting signal intensity of other tissues. Fat saturation has proven especially beneficial when applied to postcontrast T1-W images of the brain and spine as it facilitates differentiation of contrast-enhancing lesions from adjacent fat and aids in the identi-fication of meningeal enhancement (Fig. 6.20).
Magnetization transfer imaging:Magnetization transfer pulses can be applied in SE or GRE sequences to produce additional signal suppression of tissue water. The technique may be used qualitatively to increase the visibility of lesions seen during MRA and following contrast administration or quantitatively to aid in the diagnosis of white matter disease.
Artifacts37, 57, 100, 101, 132, 307, 406
A detailed discussion of MR artifacts is beyond the scope of this chapter. However, some important artifacts frequently encoun-tered when imaging the CNS in small animals will briefly be dis-cussed.
Motion:Motion artifacts are probably less common in veteri-nary than in human medicine as most patients are scanned under general anesthesia. However, even if the animal does not move during the scan, physiologically moving structures (e.g. the heart) will still result in artifacts. Motion artifacts always occur in the direction in which the phase encoding gradient was applied, regardless of the direction of motion.
They manifest as “ghosts” of the moving structure at various locations along the phase axis, blurring and/or parallel bands.
Remedies include better restraint of the patient, breath-hold techniques (limited in most systems due to duration of scan),
(a) (b)
Figure 6.20 Transverse postcontrast T1-W SE images of the canine brain without (A) and with (B) fat saturation. (A) No abnormalities are detected. (B) Suppression of adjacent bone marrow fat allows identification of meningeal enhancement, most obvious along the right occipital lobe.
cardiac/respiratory gating, presaturation pulses, flow com-pensation techniques, motion correction techniques, and flipping phase and frequency encoding gradients (if motion artifact interferes with evaluation of area of interest).
Cerebrospinal fluid (CSF) flow artifacts:A CSF flow void arti-fact appears as artificial loss of signal from CSF, which is most commonly encountered on T2-W images. It is attributed to a rapid or turbulent flow of CSF, where flowing protons move so quickly that they are not exposed to both the initial 90◦and the 180◦refocusing RF pulse. It may occasionally be seen in normal dogs, but it seems more common in small-breed dogs with increased ventricular size and syringomyelia. As this artifact is more likely to occur with a smaller slice thickness and a longer TE, modification of imaging parameters will decrease severity. However, a decrease in TE will also result in an undesirable change in weighting and may not be fea-sible. A similar artifact known as “entry slice phenomenon”
appears as an artificially high signal at a site of CSF flow when nonsaturated spins enter the imaging plane and generate a strong signal after application of the 90◦pulse (Fig. 6.21). CSF flow associated artifacts can easily be identified by compar-ison with other sequences and image planes as they will not be consistent findings.
Chemical shift artifact of the first kind:This is a misregistra-tion artifact resulting from a minimal difference in preces-sion frequency between fat and water protons. The MR unit is tuned to “listen” for hydrogen protons and expects them to precess at a certain frequency at a certain location. Hydrogen protons in fat precess ever so slightly slower than hydrogen protons in water. This minimal difference is enough to mis-place the signal from fat on the resultant image. This artifact occurs in the direction of the frequency encoding (or “read-out”) gradient and manifests at a strongly hypointense and an
Figure 6.21 CSF flow artifact. A strongly hyperintense area is observed within the mildly dilated fourth ventricle on transverse T2-W FLAIR image (arrow). This was not identified on other sequences, ruling out a true intraventricular lesion.
opposite strongly hyperintense crescent at any fluid-fat inter-face. This artifact is important to recognize as it occurs at the interface of subarachnoid space and epidural fat when imag-ing the spine and may mimic a lateralized lesion if the fre-quency encoding gradient is applied in a laterolateral (rather than dorsoventral or ventrodorsal) direction (Fig. 6.22). The effect of this artifact on diagnostic image quality can be minimized by swapping the phase and frequency encoding gradient, increasing the receiver bandwidth, and using fat-suppression techniques (see above).
Figure 6.22 Chemical shift artifact observed at the border of subarachnoid space and epidural fat in the lumbar spine of a dog. The frequency encoding gradient has been applied in a laterolateral direction. The chemical shift artifact results in a black crescent to the left and a white crescent to the right of the spinal cord (arrows) due to misregistration of the signal from fat bordering fluid. This may be confused with a lateralized lesion and is best avoided by applying the frequency encoding gradient in a dorsoventral or ventrodorsal direction.
Susceptibility:Magnetic susceptibility is a term used to describe the magnetic properties of a material. Diamagnetic materials (e.g. soft tissues) have very low susceptibility and weaken a magnetic field. Paramagnetic materials (e.g. gadolinium and certain hemoglobin degradation products) have a slightly stronger susceptibility and focally enhance a magnetic field.
Ferromagnetic materials (e.g. iron) become strongly magne-tized and experience a large force when placed in an external magnetic field. Presence of materials with differing suscep-tibility in the field of view results in suscepsuscep-tibility artifacts.
These range in severity from focal signal voids in case of hem-orrhage (see Figs 6.15, 6.38, 6.41B, 6.50) to geometric image distortion, progressive or abrupt signal void in the area of the object, and areas of sharply defined high signal intensity adja-cent to the object in case of ferromagnetic materials. Suscep-tibility artifacts may be beneficial (e.g. improved detection of small hemorrhagic infarcts) or detrimental (e.g. identifi-cation microchip immediately adjacent to area of interest).
Possible remedies include the use of (F)SE rather than GRE sequences, decreasing TE, use of STIR rather than fat satura-tion pulses, swapping phase and frequency encoding gradi-ents, and increasing receiver bandwidth.
Partial volume averaging:This artifact occurs when materials of different intensity are included in the same slice thick-ness (or even the same voxel) and their intensities are aver-aged. This artifact is significant as hyperintensities adjacent to fluid-filled structures (subarachnoid space, ventricles) on T2-W images resulting from averaging of brain and CSF sig-nal may be misinterpreted as parenchymal lesions. Remedies
include decreasing slice thickness, verification of lesions on additional planes, and verification of any T2 abnormality on FLAIR.
Contrast media in MRI62, 64, 116, 119, 150, 186, 271, 275, 276, 310, 405
MRI contrast agents most commonly used in veterinary medicine are gadolinium-based. Contrast medium is admin-istered at a dose of 0.1 mmol/kg, which may be increased to improve detection of poorly enhancing lesions. Enhancement is seen if a lesion is vascularized and is located outside the blood–
brain barrier. Gadolinium-based contrast media predominantly affect T1 relaxation, and enhancing lesions appear hyperin-tense on T1-W images (see Figs 6.6, 6.16, 6.20). Certain normal intracranial structures outside of the blood–brain barrier—such as pituitary gland, choroid plexus, trigeminal nerve, and blood vessels—show physiologic contrast uptake. Some disagreement exists as to the optimal timing of MR image acquisition follow-ing contrast medium administration (immediate vs. delayed).
However, at least for the brain, postcontrast images are usually acquired in more than one plane, and immediate and delayed images are therefore acquired inadvertently. Dynamic studies monitoring contrast enhancement (wash-in and wash-out) over time are not commonly used in veterinary medicine at this point but may be useful for the evaluation of the pituitary gland, cere-bral perfusion, and brain tumors. MR contrast agents are con-sidered very safe for use in veterinary patients, and reports of adverse effects are limited to one publication describing sus-pected anaphylactoid reactions in three dogs.
MRI guided tissue sampling49, 91, 123
MRI guided tissue sampling is rapidly gaining popularity in human medicine due to continued improvements in the field of MRI-compatible instruments and technologies. While it is not commonplace in veterinary neuroradiology at this point, initial studies have shown promising results.
MRI of the brain
MR imaging techniques94, 115, 131, 238, 286, 287
Positioning:An MR examination of the brain in small animals is typically performed in sternal recumbency. Under certain circumstances (e.g. if a subsequent examination of the cer-vical spine is to be performed), dorsal recumbency may be preferable.
Recommended standard planes and sequences:
r sagittal T2-W images
r transverse T2-W, T1-W, T2*-W and FLAIR