Download SWI Basics, Applications and Pitfalls

SWI Basics, Applications and Pitfalls
Jürgen R. Reichenbach
Medical Physics Group, Institute for Diagnostic and Interventional Radiology, University Clinics, Friedrich-Schiller-University, Jena, Germany
Susceptibility-weighted imaging (SWI) is a high spatial resolution, three-dimensional (3D), gradientecho (GRE) magnetic resonance (MR) technique [1]. It uses both the T2*-weighted magnitude and the
phase information to emphasize the magnetic susceptibility differences of various tissues or substances, such as blood products, iron, and calcification, by post-processing the magnitude images with
a phase mask [2]. Minimal intensity projection (minIP) images can further demonstrate tortuous vasculature and the continuity of vessels or abnormalities across slices [1-3]. Although most diagnostic MR
imaging ignores phase information, phase images, however, contain a wealth of information about local susceptibility changes between tissues [1-10], which can be useful in measuring iron content [4],
for instance, and other substances that change the local field.
Magnetic susceptibility is defined as the magnetic response of a substance when it is placed in an external magnetic field. The induced magnetization (owing to its susceptibility) in an object within a uniform external magnetic field distorts the uniform field outside the object. The spatial distribution of this
deviation in the external applied field is a function of the geometry of the object. For simple geometric
structures the field distributions may be obtained in analytical form (e.g., cylinders, spheres), whereas
with more complex structures, the expressions for magnetic field variations around the object are not
so straightforward. The key point, however, is that susceptibility differences between two adjacent
structures lead to a spatial field deviation within and around them, which is a function of their geometries. This field deviation, in turn, leads to local phase changes in the MR phase images. However,
phase images contain information about all magnetic fields, microscopic and macroscopic. The former
may be from small amounts of local iron deposits, for example, whereas the latter consists of field
changes caused by the geometry of the object, such as air/tissue interface effects, and by inhomogeneities in the main magnetic field. The latter tend to have a low spatial-frequency dependence (the
phase varies slowly over the image) and, for this reason, can be removed with a high-pass (HP) phase
filter. This can be achieved, for instance, with homodyne filtering [11,12].
Using phase to enhance contrast between tissues with different susceptibilities can be accomplished
in several steps. First, a high-pass filter as mentioned above can be used to remove the low-spatial
frequency components of the background field. Second, this “corrected” phase image is used to create
a “phase” mask that is used to multiply the original magnitude image to create novel contrasts in the
magnitude image [1]. The phase mask is usually applied in the following manner [1-3]: If the minimum
phase of interest is, for example –π, then the phase mask is designed to be f(x) = (ϕ(x) + π)/ π for
phases < 0, and to be unity otherwise, where ϕ(x) is the phase at location x. That is, those pixels with
a phase of –π will be completely suppressed and those with a value between –π and zero will be only
partly suppressed. This phase mask, f(x), then takes on values that lie between zero and unity and is
referred to as negative phase mask. It can be applied any number of times (integer m) to the original
magnitude image, ρ(x), to create a new image fm(x)·ρ(x) with different contrasts. Another mask might
be defined to highlight positive phase differences: ρ(x)new = gm(x) ρ(x). If the maximum phase of interest is, for example, π, then the phase mask is designed to be g(x) = (π–ϕ(x))/ π for phase > 0, and
unity otherwise. This mask is referred to as positive phase mask.
The phase mask can then be applied any number of times to the original magnitude image to create a
new image with a different contrast. After this processing, the data are often viewed by using a minimum intensity projection (mIP) over a few images (typically 4 images).
Over the years since its introduction in 1997, SWI has been found to provide additional clinically useful
information that is often complementary to conventional MR imaging sequences used in the evaluation
of various neurologic disorders, including traumatic brain injury (TBI), coagulopathic or other hemorrhagic disorders, vascular malformations, cerebral infarction, neoplasms, and neurodegenerative disorders associated with intracranial calcification or iron deposition. There exist several excellent clinical
reviews that cover many SWI applications in patients [13-21]. For instance, SWI has opened the door
for improved contrast and improved detection of hemorrhage in tumours. Part of the characterization
of tumours lies in understanding the angiographic behaviour of lesions both from the perspective of
angiogenesis and micro-hemorrhages. Since aggressive tumours tend to have rapidly growing vasculature and many microhemorrhages, the ability to detect these changes in the tumour could lead to a
better determination of its status, including better contrast in detecting tumour boundaries [14, 15, 22-
Proc. Intl. Soc. Mag. Reson. Med. 18 (2010)
24]. Multiple Sclerosis (MS) is usually studied with FLAIR and contrast enhanced T1-weighted imaging.
However, SWI reveals not only the venous connectivity in some lesions but also presents evidence of
iron in some lesions. This key new information may help understand better the physiology of MS. Recent evidence indicates that SWI has the potential to recognize the presence of iron in MS lesions,
visualize lesions missed by conventional methods and visualize different lesion characteristics [25,
26].
There are, however, a few issues that need to be taken into account when applying susceptibility
weighted imaging. The main problem encountered during the reconstruction of SWI is the fact that the
phase of an MR image is only defined in the range between 0 and 2π. When the phase exceeds 2π it
jumps back to 0. Since SWI relies on phase information, and the original phase is corrected by removing spatially slowly varying phase effects, the described data processing may also remove some of the
physiologically or pathophysiologically relevant phase information from larger anatomic structures.
Therefore, the kernel size of the homodyne filter should be restricted to 64 × 64 (if typically a 512 ×
512 matrix is acquired) to avoid adverse effects.
Furthermore, comparing phase information between subjects may be problematic because phase values depend on the geometry of the structure of interest. This may be particularly of concern if one tries
to relate phase to iron content. Measuring iron content depends on the assignment of a given phase
change in a tissue to a known amount of iron in that tissue. For instance, in the work by Haacke et al.
[13], this was done by associating the phase change in the motor cortex with 60 µg of nonheme Fe/g
tissue. This, however, ignores the fact that some of the phase shift, if not most of it, may come from
heme iron and not ferritin. In any case, one needs to be aware that the absolute relationship between
changes in phase and iron has not yet been fully established and that the relative contributions between heme and nonheme iron have not been sorted out completely.
Although high field strengths (B0 ≥ 3T) are attractive for SWI due to improved SNR, low specific absorption rates (SAR) and good phase contrast at short echo times, because the phase scales linearly
with B0·TE, one has, however, to choose appropriate sequence parameters due to the change in relaxation times with field strength and the optimum ratio between slice thickness and in-plane voxel size
for good contrast [27]. Further technical challenges include homogeneity of B0 and B1 magnetic fields,
high performance and linearity of gradient coils, and optimization of signal reception by proper design
of radiofrequency array coils. Physical problems arise from the presence of chemical shifts and large
susceptibility differences, leading to increased vulnerability towards artefacts. The advantage of increased sensitivity towards magnetic susceptibility with higher field strengths becomes particularly
critical in the vicinity to air/tissue interfaces (e.g., close to the paranasal sinuses) or bone/tissue interfaces (e.g., close to the petrous portion of the temporal bone) that result in severe static macroscopic
B0 field inhomogeneities. These macroscopic inhomogeneities lead to enhanced intravoxel signal
dephasing and cause steep topologies of the signal phase.
Another issue, which one has to consider with respect to venous contrast of SWI scans relates to examination of patients under anesthesia. Since anesthetics affect respiration and the cardiovascular
system, it is quite likely that a correlation exists between general anesthesia and venous contrast in
SWI. First indications on attenuation of signal intensity on SWI and poor demonstration of the venous
anatomy were reported for pediatric patients under anesthesia (2%–3% sevoflurane in oxygen) [28].
The reason for the loss of contrast is that sevoflurane slightly decreases the regional oxygen extraction fraction (rOEF) compared to awake patients. Since the attenuation of venous vessels varied
among anesthetic patients, it was hypothesized that the depth of anesthesia might be responsible for
this variation. To prove this hypothesis, Sedlacik et al. monitored blood pressure and end tidal CO2 to
assess the depth of anesthesia (anesthetic: propofol, 150-300 mcg/kg/min) in 108 SWI pediatric examinations [29]. Based on the significantly lower venous contrast in case of reduced blood pressure
and increased end tidal CO2 and vice versa, the authors concluded that the attenuation of veins in
SWI depends indeed on the depth of anesthesia. Due to the correlation between depth of anesthesia
and CBF, the source for the venous contrast variations is most likely due to CBF changes induced by
anesthesia. However, due to the different characteristics of anesthetic agents further studies are required to improve the understanding of the relationship of anesthetic agent, concentration, and CBF
on venous contrast in SWI. Along similar lines, exceptional flow conditions should be kept in mind
when interpreting SW images in patients with unusual findings [30].
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