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Ultrafast MR imaging of the pelvis
European Journal of Radiology 29 (1999) 233 – 244
Ultrafast MR imaging of the pelvis Eric K. Outwater * Department of Radiology, Thomas Jefferson Uni6ersity Hospital, 132 South Tenth Street, 1096 Main, Philadelphia, PA 19107 -5244, USA Received 27 November 1998; accepted 30 November 1998
Abstract MR gradient systems with higher slew rates and gradient amplitude enable certain forms of imaging that are not practical with older gradient systems. These newer pulse sequences include single shot half-Fourier T2-weighted images and echo planar imaging. More important in MR imaging of the pelvis, these gradient systems benefit more conventional imaging methods such as gadolinium-enhanced 3D MR angiography, dynamic gradient echo contrast-enhanced images, and T2-weighted fast spin echo images, by shortening echo times. For most MR imaging of the pelvis, spatial resolution is paramount, and therefore sequences such as half-Fourier acquisition Turbo spin echo (HASTE) and 3D gadolinium-enhanced dynamic imaging play a less important role than in the upper abdomen. The potential of these techniques for diffusion or perfusion studies in the pelvis has not been explored. © 1999 Elsevier Science Ireland Ltd. All rights reserved. Keywords: MR imaging; Pelvis; Signal-to-noise; Fast spin echo
1. Introduction The development of fast imaging sequence techniques has greatly enhanced clinical MR imaging of the pelvis. MR imaging in the pelvis is often limited by motion artifacts. Fast imaging can alleviate or circumvent motion effects mainly due to respiratory motion of the anterior pelvic wall, bowel peristalsis, gross fetal movements, and vascular pulsations. For imaging the inner pelvis, including the prostate, bladder, vagina, normal sized uterus, and ovaries, respiratory motion is not usually significant. Therefore, imaging in suspended respiration is not advantageous and more standard techniques with better resolution and signal-to-noise (SNR) give better results. However, if the structures of interest are displaced superiorly, by a filled bladder, enlarged uterus, or pelvic mass, then respiratory motion becomes a significant component of image degradation. This means that the benefits of fast
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imaging become greater in the upper pelvis or when the uterus is enlarged. Bowel peristalsis causes motion artifacts which affect MR images of the abdomen and pelvis in two ways. First, the incoherent motion causes ghosting artifacts to propagate across the image, leading to image degradation of structures around the bowel, such as the ovaries and retroperitoneum. Second, motion of the bowel causes blurring and ghosting of the bowel loops themselves, so that detection of intrinsic bowel abnormalities is compromised. One of the major limitations of MR imaging compared to CT is its relative lack of sensitivity to bowel abnormalities. Three approaches to minimizing bowel motion artifacts are used. Firstly the patient can be imaged in the fasting state to minimize bowel motion. Secondly, an antiperistaltic agent can be given to suppress bowel motion. Lastly, the pelvis can be imaged with fast sequences to minimize motion during image acquisition. Single shot fast spin echo or half-Fourier acquisition Turbo spin echo (HASTE) type T2-weighted images are particularly adept at imaging the bowel with a minimum of motion artifact [1–3].
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Overcoming the effects of respiratory, bowel, and fetal motion on MR images has been particularly important for T2-weighted image because of the longer acquisition time required. This review will explore some of the ways ultrafast imaging can be used in the pelvis.
2. Fast T2-weighted MR imaging of the pelvis Rapid T2-weighted imaging of the abdomen and pelvis has been a major goal of pulse sequence development in magnetic resonance imaging (MRI). Specifically, high quality breath-hold T2-weighted MR sequences have been considered critical for imaging the liver and upper abdomen. The development of fast spin echo images for T2-weighting has been critical for both abdominal and pelvic imaging. For the upper abdomen, however, fast spin echo images are still prone to considerable motion artifact from movement of the diaphragm and anterior abdominal wall. This necessitated developing such techniques as respiratory triggering and the development of even faster sequences such as single shot fast spin echo. In the pelvis, however, fast spin echo images generally produce high quality T2-weighted images. Motion artifact was restricted to the upper pelvis or involved structures that were not of critical concern, such as the bowel. For this reason, regular fast spin echo sequences accomplish most of the tasks required of T2-weighted image sequences in the pelvis. Faster images may be desirable, but sacrifices in SNR or image blurring currently limits their ability to substitute for a conventional fast spin echo images.
2.1. Fast spin echo imaging in the pel6is Fast spin echo in the pelvis proved to be a major advance over conventional T2-weighted spin echo imagining for several reasons. Firstly, there is a time savings in the acquisition of multiple phase encoding steps in one TR interval. Therefore, unlike conventional spin echo images, fast spin echo images can be acquired in a fairly short time interval, on the order of 3–6 min (Fig. 1). This time savings can be converted into: (a) higher order matrices (512× 512 or 256×256 rather than the 256× 128 or 256× 192 than are used with conventional spin echo images); (b) T2-weighted images in multiple planes which are important for imaging the pelvis in general; or (c) thinner sections requiring more slices which achieves greater resolution in the slice direction. Secondly, because of the length of the echo train, fewer slices were attainable in a multislice format with fast spin echo compared to conventional spin echo images for a given TR. The easiest way to circumvent this is to use a much longer TR for fast spin echo sequence. This longer TR has the additional advantage
of removing unwanted T1 effects from the image. Lastly, fast spin echo imaging affords the opportunity to image with as long a TE as desired without paying a penalty in terms of increased motion artifact or fewer slices per TR in a multislice mode. For all these reasons, fast spin echo entirely superseded conventional spin echo for T2-weighted imaging in the pelvis. The details of fast spin echo imaging as a pulse sequence has been detailed elsewhere and will not be elaborated upon here [4–7]. Some important differences exist with fast spin echo imaging compared to conventional spin echo imaging that are important to bear in mind (Fig. 1). These differences become particularly pertinent when comparing fast spin echo sequences to single shot techniques such as HASTE, because certain adverse effects of spin echo imaging and fast spin echo imaging become even more limiting in the HASTE type sequence. The most important of these effects is image blurring due to discontinuities in k-space from phase encoding step to phase encoding step [5–7]. In conventional spin echo imaging, all phase encoding steps have the same actual echo time (TE) so that differences in overall signal intensity of a signal (a line in k-space) are due to intravoxel phase dispersion resulting from the phase encoding gradients. In fast spin echo imaging this is not the case. Each phase encoded signal in a given echo train has a different TE. Therefore, overall signal intensity of a signal in k-space is due to not only to the phase encoding gradients but to the actual TE of that signal. These discontinuities give rise to image blurring after Fourier transformation of the phase encoded signals (Fig. 1) [7]. Because the difference in signal intensity from phase-encoding step to phase-encoding step will be greater with decreasing T2 of a structure, this blurring effect is most evident in shorter T2 structures (Fig. 2) [7]. Conversely, structures with long T2 such as those containing fluid will suffer little loss in signal intensity from phase encoding step to phase encoding step and therefore, no blurring of fluid-containing structures occurs. The blurring effect becomes more acute with longer echo train lengths such as those used for single-shot fast spin echo (HASTE) sequences. Another effect of fast spin echo sequences is magnetization transfer. Magnetization transfer results from radiofrequency (RF) pulses applied off-resonance relative to a given slice. That is, when imaging several slices in multislice format, excitation of adjacent slices will affect the longitudinal magnetization of a given reference slice. This effect occurs from saturation of water protons associated with macromolecules, particularly those hydrogen-bonded to macromolecules. The degree of magnetization transfer is greater with fast spin echo sequences than conventional spin echo sequences although it occurs to some degree in both [8]. It may be more acute due to the very rapid deployment of 180° pulses in HASTE type sequences.
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Fig. 1. FSE sequences showing basic tissue contrast in the pelvis. (A) T2-weighted conventional spin echo image (TR/TE= 2350/90) in the sagittal plane through the midline shows the normal zonal anatomy of the uretus with, hyperintense endometrium, low signal intensity junctional zone, and higher signal intensity outer myometrium. (B) T2-weighted sagittal FSE sequence (TR/TEeff =3000/90) shows the same basic tissue contrast as in (A) but was performed in 3 min and 24 s. (C) T2-weighted sagittal single shot FSE sequence (TR/TEeff =25338/98) was performed in 25 s for an equivalent number of slices. Note that the signal intensity basic tissue contrast is the same as in (B) but there is slight blurring of the soft tissues of the pelvis. However, note the sharp delineation of the fluid in the bladder, vaginal fornix and endocervical canal (arrowheads).
The development of high performance gradient systems has improved fast spin echo sequences. Deployment of faster gradients has enabled shorter echo spacing, that is the time between the 180° pulses, and thus the time between the signals. This has several beneficial effects. Firstly, reducing the echo spacing reduces the time allowed for T2 decay from phase encoding step to phase encoding step. This effectively reduces image blurring and shortens the overall time it takes to deploy the echo train. Decreasing the echo spacing also has the effect of diminishing motion artifacts [9]. Lastly, improved gradient systems make acquisition of longer echo trains, such as those used in single shot techniques described in the next section, more practical.
2.2. Single shot fast spin echo HASTE images were the first type of sequence to obtain high quality T2-weighted spin echo-type images in a single breathhold. Previous attempts to obtain T2-weighted images in a breathhold included T2-weighted gradient echo images such as steady state free precession. These suffered from low resolution and considerable magnetic susceptibility artifacts that limited their use in the abdomen and pelvis. HASTE sequences employ half-Fourier acquisition and a long echo train fast spin echo sequence to obtain all the data necessary to form k-space from a single acquisition or a single ‘shot’. Therefore, these sequences are referred to as HASTE or single shot fast spin echo.
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Fig. 2. Fast spin echo compared to single shot fast spin echo in a woman with uterine myomas. (A) Fast spin echo image sequence (TR/TEeff = 3233/90) through the uterus shows muliple uterine myomas of varying signal intensity. (B) Single shot FSE sequence (TR/TEeff = 17137/98) shows lower signal-to-noise (SNR) overall. There is blurring of the tissues of the uterus and myomas as well, but the urine in the bladder remains sharp.
Single shot fast spin echo sequences can be used to obtain relatively motion free breathold sequences in the pelvis with tissue contrast similar to more conventional fast spin echo. For conventional pelvic imaging, the advantage of these sequences is the ability to image the abdominal wall, mesenteric fat and bowel without motion. The main disadvantage of these sequences is the overall lower SNR and the tissue blurring in the phaseencoding direction due to alterations to the pointspread function, as previously described for fast spin echo sequences. These effects are more pronounced on single shot fast spin echo sequences and are the major factor limiting their more widespread use in the abdomen and pelvis (Fig. 2). Single shot fast spin echo sequences have been substituted for fast spin echo sequences for imaging the pelvis (Fig. 2). For certain specific, ‘simple’ applications such as identification of fibroid, single shot fast spin echo sequences may be able to obtain a study in a shorter period of time [10]. Imaging of the bowel is superior with single shot fast spin echo sequences, for two reasons [3]. Imaging time is sufficiently short so that perstalsis-related motion artifacts are reduced. Also, the long T2 of the luminal contents will cause the mucosal surfaces to be imaged with high resolution.
2.3. Echo planar imaging Echo planar imaging has not been widely used for imaging the pelvis. The reasons for this are several fold. First, until recently, gradient performance has not been
up to the task of the rapid gradient switching that is necessary to generate echo planar images on standard MR imaging systems. Secondly, magnetic susceptibility and chemical shift artifacts are a problem with echoplanar imaging and would greatly limit its potential in the pelvis. Magnetic susceptibility is high at interfaces between gas and soft tissues. This situation occurs commonly in the tortuous sigmoid colon, small bowel, and rectum and therefore, structures adjacent to these may be distorted or obscured on echo planar images. Lastly, it has not been considered necessary to obtain the degree of temporal resolution that echo planar imaging offers. In general, for imaging the pelvis, spatial resolution is at a premium. Contrast resolution or tissue contrast is less important and temporal resolution may be the least important of all. Some degree of temporal resolution is important when studying dynamics of contrast enhancement and vascularity, but this type of information has rather restricted utility in the pelvis.
2.4. Obstetric imaging MR images of the intrauterine fetus have always been susceptible to motion induced artifact [11,12]. This motion has two sources. One is respiratory induced movement of the uterus due to maternal or respiratory motion. The second is gross movements of the fetus itself. Motion of the fetus causes motion of the amniotic fluid around it, which generates a marked ghosting artifact and blurring such that, in order to image the fetus for possible fetal abnormalities, paralysis of the
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Fig. 3. Single shot FSE imaging of gravid uterus. (A) Single shot FSE sequence (TR/TE 19062/98) shows excellent anatomic detail in the fetus. Note the identification of the liver (li) lungs (lu) and the severe hydraencephaly (h). The low signal intensity vessels of the umbilical cord lie on either side of the fetus’s neck. (B) Axial single shot FSE sequence with the same parameters as (A) shows hydraencephaly with some preservation in a vascular territory of the occipital lobe (arrow). Note the delineation of the placenta posteriorly (p).
fetus was frequently necessary with conventional spin echo sequences or even fast spin echo sequences. Fast spin echo sequences are suitable for imaging the placenta and uterus but frequently give poor quality images of the fetus itself due to motion induced artifact. The single shot sequences (HASTE) have dramatically improved the potential for MR imaging of the obstetric patient. These sequences are obtained in a breathhold but this is not nearly as important as the overall speed of imaging. What is more critical is that these images be acquired quickly before gross fetal motion has occurs. Frequently some of the images will be degraded if motion occurs during their acquisition; whereas others will be relatively free of motion artifact when acquired in the absence of fetal motion. Therefore, repeated acquisition of the sequences can be performed to obtain some diagnostic images that are free of motion artifact. The blurring of soft tissue structures demonstrated in HASTE sequences is of lesser importance when imaging the fetus because most structures are bordered by fluid, and therefore image sharply. Examining the fetus with HASTE imaging has been made easier by obtaining oblique planes that represent anatomic cross sections realtive to the fetus, not the mother (Fig. 3). Single shot fast spin echo sequences are useful for imaging the fetus or placenta in cases where ultrasound findings are equivocal. The display of fetal anatomy before 20 weeks of gestational age is usually inferior to that of ultrasound, due to motion and the small size of the fetus [13,14]. Single shot fast spin echo images show
particular advantage compared to ultrasound in situations where ultrasound is limited: in the setting of oligohydramnios and in late third trimester. MR imaging can demonstrate additional findings in fetuses with neurologic abnormalities, such as germinal matrix hemorrhages and agenesis of the corpus callosum [15–19]. MR imaging is also useful in showing the contents of diaphragmatic hernias and omphaloceles because liver can easily be distinguished from bowel [20]. Widespread use of the single shot fast spin echo sequence for fetal imaging awaits research into the optimal indications for MR imaging in this population. Another approach to imaging the fetus uses echo planar imaging [21]. While echo planar imaging can easily achieve images in less than a second like single shot fast spin echo, SNR considerations probably favor the latter. Echo planar imaging has been used to investigate aspects of lung and liver maturation [22,23].
2.5. MR urography T2-weighted fast spin images can be performed with a sufficiently long TE to effectively image only fluid [24]. This sequence is termed MR urography when applied to the pelvis and MR cholangiography when applied to the biliary tree. The basic principle is the same: to acquire an image or images which display the entire fluid-filled collecting system. This can be accomplished by one image representing a slab thick enough to include the renal pelvis, ureters and bladder [25,26]. An alternative is to acquire contiguous thin-section images and perform a computer generated reconstruc-
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Fig. 4. FSE sequences used for MR urography. (A) Axial T2-weighted FSE image through the base of the bladder (TR/TEeff =4350/119) shows a tumor invading through the muscularis of the bladder. (B) MR urogram made from 3 mm thick coronal sections using a FSE sequence (TR/TEeff = 12 000/260, 4 NEX) shows bilateral ureters. The right ureter is dilated down to the level of the tumor mass, but the site of the obstruction (arrow) is not clearly delineated on this anterior to posterior projection. Note the right hydronephrosis. (C) Cropped view of the distal right ureter shows the dilated ureter (arrowheads) terminating at the level of the tumor at the ureterovesicle junction (arrow). This is a lateral projection view of a urogram made sagittal images.
tion (e.g. maximum intensity projection or MIP) from multiple perspectives to visualize the urinary system (Fig. 4). The single section slab approach has the advantage of fast acquisition, which allows acquisition in suspended respiration [25,26]. MR urograms can show the site of urinary tract obstruction very well but because the unobstructed ureters are intermittently peristalsing, they are generally not shown in their entirety. Unfortunately image resolution is not equivalent to intravenous urograms or retro-
grade pyelograms, so they cannot be used as a substitute for these in many applications. In particular, the detection of transitional cell neoplasia and small calculi may not be optimal on these images.
3. Ultrafast vascular imaging in the pelvis Vascular imaging of the pelvis has always been a major challenge for body vascular imaging using conventional
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Fig. 5. A 71-year-old man with femoral–femoral cross graft. Time-of-flight (TOF) imaging compared to 3D gadolinium enhanced imaging in the pelvis. (A) Anterior to posterior projection reconstruction of individual axial 2D TOF gradient echo sequences (TR/TE 33/6.4), flip angle of 60° shows the graft entering the right common femoral artery and bilateral occulsion of the native iliac arteries. The femoral – femoral cross graft is not well shown due to inplane flow saturation. (B) 3D gadolinium enhanced gradient echo acquisition (TR/TE 11.2/2.1) flip angle of 60° shows greater anatomical detail in the arteries compared to (A). Note that the femoral – femoral cross graft (arrowheads) is well shown with the 3D gadolinium enhanced technique. Also note that the left external iliac artery (arrow) is shown in (B) but is not shown in (A). This is due to reversal of flow with filling via the cross graft.
techniques. Of the various MRA sequences available, time-of-flight (TOF) gradient echo imaging was used most widely prior to the development of faster imaging methods. TOF imaging generally works well for imaging the pelvic veins, whereas the pelvic arteries have provided a greater challenge [27]. This is partly due to the fact that prominent arterial pulsations are normally present in the pelvic arteries, leading to marked differences in flow-related enhancement from phase encoding step to phase encoding step. The changes in flow-related enhancement leads to prominent ghosting artifacts and a resultant loss of signal intensity in the pelvic arteries. Additionally, the pelvic arteries are frequently tortuous and thus susceptible to in plane flow saturation on TOF images (Fig. 5). Lastly, the pelvic arteries were vulnerable to magnetic susceptibility artifact resulting from neighboring colon. Because the pelvic arteries are of critical imaging importance in evaluating patients with claudication or tissue loss, precise delineation of any stenoses is crucial. For these reasons, optimal imaging of pelvic arteries was not possible before the development of fast imaging techniques with gadolinium enhancement [27]. 3D MRA with gadolinium proved to be a major advance in overcoming the special problems posed by pelvic arteries [28]. With the short echo times used in 3D MRA, magnetic susceptibility artifacts were no longer a significant source of error. The dramatically shortened T1 of blood provided by gadolinium eliminates the in plane flow saturation that was frequently
seen in torturous arteries (Fig. 5). Lastly, because the 3D MRA technique relies on imaging blood by virtue of its shortened T1 and not flow-related enhancement per se, this technique is not susceptible to the alterations in flow-related enhancement which led to the view-to-view errors and pulsation artifacts of the 2D TOF technique [28]. The details of the gadolinium enhanced 3D gradient echo imaging technique have been described elsewhere [28]. Briefly, this technique involves a 3D gradient echo acquisition to image a volume with thin sections amenable to 3D reconstruction and reformatting. This type of sequence can be performed on more conventional gradients systems; however, the high performance gradients with higher gradient strengths and slew rates allow the 3D sequence to be performed in a shorter time period. Acquisition of the 3D sequence during a shorter time period means that timing the acquisition to the bolus is less problematic, and the sequence can be performed with breathholding to eliminate some sources of motion. For imaging a given volume, an orientation to the slab must be chosen. This differs from TOF MRA, where the pelvis is imaged in cross-section (axially) and final reconstructions are made in the anterior–posterior projection. Thus, with TOF MRA final resolution depends on slice thickness. For imaging the pelvis with a 3D volume acquisition, the coronal plane or slab includes the target vessels, and the section thickness determines the final resolution. The 3D gradient echo
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Fig. 6. Acquisition of more than one phase in the 3D gadolinium enhanced MR angiography sequence. A 73-year-old woman with infected femoral graft and venous thrombosis. (a) 3D gadolinium enhanced projection reconstruction of a coronal slab from a gradient echo sequence (TR/TE 13/2.3) shows a completely occluded iliac arterial system. The right arterial system is severely diseased with multiple stenosis. A cross graft femoral to popliteal graft is seen (arrowheads). (B) The venous phase of the 3D gadolinium enhanced technique shows enhancement around the graft focally (arrowheads) which prove to represent infection in the graft. Note the numerous venous collaterals from bilateral iliac deep venous thrombosis.
sequence is obtained with minimum TE and TR. Because the large number of phase encoded steps involved in the acquisition, the minimum TR obtainable will dramatically effect the total acquisition time. Gadolinium is injected intravenously and the plateau of arterial phase enhancement is timed to correspond to the center of k-space acquisition of the 3D gradient echo scan. If the 3D sequence is a half-Fourier acquisition, then the center of a k-space may occur at the beginning of the acquisition rather than in the middle. 3D gadolinium enhanced sequences are less susceptible to artifacts in general than 2D TOF imaging. Several studies have shown greater accuracy with the 3D gadolinium enhanced technique compared to the 2D TOF technique [29 – 33]. Pitfalls in the 3D gadolinium enhanced technique are primarily related to technical problems with the performance of the examination. Because gadolinium dramatically shortens the T1 of the blood, selective in-flow saturation of venous vessels is far less effective with the gadolinium enhanced technique compared to the TOF technique. This is compounded by the fact that the usual acquisition for the gadolinium enhanced sequence is in the coronal plane. Therefore, saturation bands for suppression of venous in-flow is not possible, and selective imaging of the arterial systems depends upon critical timing so that venous enhancement does not occur. A corollary to this is if the sequence is rapidly repeated after the arterial
phase then a venous phase ‘venogram’ may be obtained (Fig. 6). 3D gadolinium enhanced MRA is particularly well suited to patients with complex arterial grafts in place. Imaging of femoral bypass grafts was a particular problem with the 2D TOF imaging due to in-plane flow saturation. With the 3D gadolinium enhanced technique, the flow direction is irrelevant and therefore, flow in grafts running in any direction can be effectively imaged (Fig. 5). Surgical clips may still cause artifactual stenosis with the 3D gadolinium enhanced sequence, but the problem will be less profound compared with 2D TOF imaging because the TE of the 3D sequence is so much shorter. 3D gadolinium enhanced sequences have been used in a number of applications in the pelvis [33,34]. They are more accurate for identification of stenoses in the iliac arteries than conventional TOF sequences. They are useful for identification of stenosis and thrombosis affecting renal transplants, arteries and veins [35]. Because of the short transient time of arterial blood through the kidneys, it is relatively simple to image both the arteries and veins with this type of sequence. These sequences can detect arterial lesions associated with impotence [36]. The potential of these sequences for delineation of arterial supply in other pelvic disorders per se has not been explored.
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Fig. 7. Dynamic gadolinium enhancement of the uterus in a woman with cervical carcinoma. Sagittal T2-weighted fast spin echo image (TR/TE= 3883/126) shows infiltration of the cervix (arrow) with slightly higher than normal tissue signal intensity. (A) Dynamic gradient echo multiplanar images (TR/TE= 125/2.5) performed before (B), in the arterial phase (C), and venous phase of gadopentatate dimeglumine enhancement. The cervix, which normally enhances less than uterine corpus, is enhancing more than corpus myometrium, indicating diffuse infiltration with tumor.
4. Fast gadolinium-enhanced imaging in the pelvis Dynamic imaging with contrast enhancement is a valuable tool in several areas of the body, most prominently breast, liver and pancreas. Dynamic imaging of the pelvis has been explored to some extent and is primarily useful in the differentiation of scar versus recurrent tumor in the postoperative pelvis and for evaluation of endometrial abnormalities and bladder tumors (Fig. 7) [37–44]. It is probably not useful in evaluation of prostatic disorders or adnexal disorders, although it has not been thoroughly explored for these applications.
For the purpose of dynamic images, three parameters are frequently in conflict with one another in providing the maximal information possible. These are: temporal resolution (i.e. the number of images of a given area obtained per unit time), spatial resolution, and slice coverage (or the volume of tissue that is scanned). The best temporal resolution from low-resolution images are obtained at one slice position, while repeating the image rapidly during contrast administration. This sequence has been used for imaging prostatic lesions, individual lymph nodes, and cervical and bladder tumors (Fig. 7). The problem with this approach is that it gives information about only one plane rather than the
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Fig. 8. Three-dimensional gadolinium-enhanced imaging of the pelvis in a 54-year-old woman with a cystic mass of the left ovary. (A) Sagittal T2-weighted fast spin echo image (TR/TEeff = 3250/105) shows a multilocular cyst of the left ovary (arrow). (B) Sagittal fat-suppressed 3D gradient echo image (TR/TE=7.0/2.3 ms) is performed with 32 4 mm thick sections and a 15° flip angle in 26 s. Note the sharp delineation of the bowel (white arrow). (C) Same sequence after gadopentatate dimeglumine administration shows enhancement of the cyst walls (arrow) and pelvic vessels. (D) For comparison, a 2D gradient echo image (TR/TE= 400/2.1) yields higher signal-to-noise (SNR) and resolution but is not performed with suspended respiration. Local sigmoid wall thickening from Crohn’s disease is fistulized (thin arrow) to the left ovary (black arrow).
pelvis or tumor as a whole. Another approach to dynamic imaging involves using higher resolution images with a longer TR to allow multiplanar imaging of much if not all of the pelvis. However, the temporal resolution of this type of imaging is much lower. Generally, images corresponding to arterial phase, venous phase, and parenchymal or interstitial phases are obtained. With faster and more powerful gradients, dynamic imaging with 3D T1-weighted acquisition becomes possible. This type of sequence can be used for dynamic
gadolinium enhancement of organs in the arterial, venous, and parenchymal phases. There are two reasons for performing 3D acquisitions for dynamic scanning. Firstly, with similar parameters, 3D acquisitions yield higher SNR acquisitions. Partly for this reason, thinner sections can be obtained without cross talk, providing higher spatial resolution in the slice select direction (Fig. 8). Secondly, 3D acquisitions can be reconstructed to provide MR angiograms in the arterial phase since these sequences are fundamentally similar to the 3D gadolinium enhanced MR angiography sequences.
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There are some disadvantages to using the 3D gadolinium enhanced technique for dynamic imaging of the pelvis: (1) there is competition between the parameters necessary for high resolution imaging of the pelvis, e.g. a 20 cm field-of-view may not be implemented with dynamic images due to unacceptable demands on the gradients necessary to obtain the proper TE. Secondly, the 3D sequences use a very short TE and TR. The short TR lends very little T1-weighted contrast to the image (Fig. 8). Thus, relative advantages and disadvantages must be considered for a given clinical application before showing 3D gadolinium enhances images over 2D gradient echo images.
5. New directions The potential of newer techniques, such as perfusion or diffusion imaging in MR imaging of the pelvis, has been largely unexplored. In the abdomen, measuring the apparent diffusion coefficient (ADC) may help discriminate hemanigomas from solid tumors [45 – 47]. Chronic renal failure may be associated with alterations in the ADC as well [48]. However, little is known about the state of water diffusion in pelvic organs. Yang et al. found that in rats water diffusion is anisotropic in the myometrium, but not in the endometrium [49]. Apparent water diffusion coefficients depend on hormonal influences in the endometrium, but not in the myometrium. These preliminary animal studies suggest that diffusion imaging of the human uterus may be useful in studying endometrial physiology [49]. Fast spin echo sequences can be modified with diffusion-sensitizing gradients [50]. These sequences can also be used to study perfusion with arterial spin labeling [51]. Practical and convenient fast sequences such as these may spur research into the practical applications of diffusion and perfusion imaging in the pelvis. Ultrafast imaging is important in real-time MR ‘flouroscopic’ techniques that may prove useful for interventional procedures [52,53]. The clinical potential of MR interventional procedures in the pelvis has not been explored. Possible applications include interventional thermal ablation of tumors, biopsies, and radiation seed placement [54].
6. Summary Ultrafast imaging is made possible by stronger gradients capable of generating higher slew rates and thus much more rapid deployment of gradients. These gradients on conventional MR systems (Siemens Vision, GE EchoSpeed, Picker EDGE, Philips ACSNT) accelerate and improve pulse sequences that are important in imaging the pelvis, such as fast spin echo, fast spoiled
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gradient echo, and 3D fast gradient echo. Other sequences such as single shot fast spin echo (HASTE) and echo planar suffer from lower spatial resolution and certain artifacts which limit their effectiveness in the pelvis, where spatial resolution is critical. These sequences could be useful for diffusion imaging, the potential of which is largely unexplored in the abdomen. 3D gadolinium-enhanced MRA techniques are generaly superior for imaging the pelvic arteries.
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Ultrafast MR imaging of the pelvis Eric K. Outwater * Department of Radiology, Thomas Jefferson Uni6ersity Hospital, 132 South Tenth Street, 1096 Main, Philadelphia, PA 19107 -5244, USA Received 27 November 1998; accepted 30 November 1998
Abstract MR gradient systems with higher slew rates and gradient amplitude enable certain forms of imaging that are not practical with older gradient systems. These newer pulse sequences include single shot half-Fourier T2-weighted images and echo planar imaging. More important in MR imaging of the pelvis, these gradient systems benefit more conventional imaging methods such as gadolinium-enhanced 3D MR angiography, dynamic gradient echo contrast-enhanced images, and T2-weighted fast spin echo images, by shortening echo times. For most MR imaging of the pelvis, spatial resolution is paramount, and therefore sequences such as half-Fourier acquisition Turbo spin echo (HASTE) and 3D gadolinium-enhanced dynamic imaging play a less important role than in the upper abdomen. The potential of these techniques for diffusion or perfusion studies in the pelvis has not been explored. © 1999 Elsevier Science Ireland Ltd. All rights reserved. Keywords: MR imaging; Pelvis; Signal-to-noise; Fast spin echo
1. Introduction The development of fast imaging sequence techniques has greatly enhanced clinical MR imaging of the pelvis. MR imaging in the pelvis is often limited by motion artifacts. Fast imaging can alleviate or circumvent motion effects mainly due to respiratory motion of the anterior pelvic wall, bowel peristalsis, gross fetal movements, and vascular pulsations. For imaging the inner pelvis, including the prostate, bladder, vagina, normal sized uterus, and ovaries, respiratory motion is not usually significant. Therefore, imaging in suspended respiration is not advantageous and more standard techniques with better resolution and signal-to-noise (SNR) give better results. However, if the structures of interest are displaced superiorly, by a filled bladder, enlarged uterus, or pelvic mass, then respiratory motion becomes a significant component of image degradation. This means that the benefits of fast
* Tel.: +1-215-955-1535; fax: +1-215-955-5329. E-mail address: [email protected] (E.K. Outwater)
imaging become greater in the upper pelvis or when the uterus is enlarged. Bowel peristalsis causes motion artifacts which affect MR images of the abdomen and pelvis in two ways. First, the incoherent motion causes ghosting artifacts to propagate across the image, leading to image degradation of structures around the bowel, such as the ovaries and retroperitoneum. Second, motion of the bowel causes blurring and ghosting of the bowel loops themselves, so that detection of intrinsic bowel abnormalities is compromised. One of the major limitations of MR imaging compared to CT is its relative lack of sensitivity to bowel abnormalities. Three approaches to minimizing bowel motion artifacts are used. Firstly the patient can be imaged in the fasting state to minimize bowel motion. Secondly, an antiperistaltic agent can be given to suppress bowel motion. Lastly, the pelvis can be imaged with fast sequences to minimize motion during image acquisition. Single shot fast spin echo or half-Fourier acquisition Turbo spin echo (HASTE) type T2-weighted images are particularly adept at imaging the bowel with a minimum of motion artifact [1–3].
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Overcoming the effects of respiratory, bowel, and fetal motion on MR images has been particularly important for T2-weighted image because of the longer acquisition time required. This review will explore some of the ways ultrafast imaging can be used in the pelvis.
2. Fast T2-weighted MR imaging of the pelvis Rapid T2-weighted imaging of the abdomen and pelvis has been a major goal of pulse sequence development in magnetic resonance imaging (MRI). Specifically, high quality breath-hold T2-weighted MR sequences have been considered critical for imaging the liver and upper abdomen. The development of fast spin echo images for T2-weighting has been critical for both abdominal and pelvic imaging. For the upper abdomen, however, fast spin echo images are still prone to considerable motion artifact from movement of the diaphragm and anterior abdominal wall. This necessitated developing such techniques as respiratory triggering and the development of even faster sequences such as single shot fast spin echo. In the pelvis, however, fast spin echo images generally produce high quality T2-weighted images. Motion artifact was restricted to the upper pelvis or involved structures that were not of critical concern, such as the bowel. For this reason, regular fast spin echo sequences accomplish most of the tasks required of T2-weighted image sequences in the pelvis. Faster images may be desirable, but sacrifices in SNR or image blurring currently limits their ability to substitute for a conventional fast spin echo images.
2.1. Fast spin echo imaging in the pel6is Fast spin echo in the pelvis proved to be a major advance over conventional T2-weighted spin echo imagining for several reasons. Firstly, there is a time savings in the acquisition of multiple phase encoding steps in one TR interval. Therefore, unlike conventional spin echo images, fast spin echo images can be acquired in a fairly short time interval, on the order of 3–6 min (Fig. 1). This time savings can be converted into: (a) higher order matrices (512× 512 or 256×256 rather than the 256× 128 or 256× 192 than are used with conventional spin echo images); (b) T2-weighted images in multiple planes which are important for imaging the pelvis in general; or (c) thinner sections requiring more slices which achieves greater resolution in the slice direction. Secondly, because of the length of the echo train, fewer slices were attainable in a multislice format with fast spin echo compared to conventional spin echo images for a given TR. The easiest way to circumvent this is to use a much longer TR for fast spin echo sequence. This longer TR has the additional advantage
of removing unwanted T1 effects from the image. Lastly, fast spin echo imaging affords the opportunity to image with as long a TE as desired without paying a penalty in terms of increased motion artifact or fewer slices per TR in a multislice mode. For all these reasons, fast spin echo entirely superseded conventional spin echo for T2-weighted imaging in the pelvis. The details of fast spin echo imaging as a pulse sequence has been detailed elsewhere and will not be elaborated upon here [4–7]. Some important differences exist with fast spin echo imaging compared to conventional spin echo imaging that are important to bear in mind (Fig. 1). These differences become particularly pertinent when comparing fast spin echo sequences to single shot techniques such as HASTE, because certain adverse effects of spin echo imaging and fast spin echo imaging become even more limiting in the HASTE type sequence. The most important of these effects is image blurring due to discontinuities in k-space from phase encoding step to phase encoding step [5–7]. In conventional spin echo imaging, all phase encoding steps have the same actual echo time (TE) so that differences in overall signal intensity of a signal (a line in k-space) are due to intravoxel phase dispersion resulting from the phase encoding gradients. In fast spin echo imaging this is not the case. Each phase encoded signal in a given echo train has a different TE. Therefore, overall signal intensity of a signal in k-space is due to not only to the phase encoding gradients but to the actual TE of that signal. These discontinuities give rise to image blurring after Fourier transformation of the phase encoded signals (Fig. 1) [7]. Because the difference in signal intensity from phase-encoding step to phase-encoding step will be greater with decreasing T2 of a structure, this blurring effect is most evident in shorter T2 structures (Fig. 2) [7]. Conversely, structures with long T2 such as those containing fluid will suffer little loss in signal intensity from phase encoding step to phase encoding step and therefore, no blurring of fluid-containing structures occurs. The blurring effect becomes more acute with longer echo train lengths such as those used for single-shot fast spin echo (HASTE) sequences. Another effect of fast spin echo sequences is magnetization transfer. Magnetization transfer results from radiofrequency (RF) pulses applied off-resonance relative to a given slice. That is, when imaging several slices in multislice format, excitation of adjacent slices will affect the longitudinal magnetization of a given reference slice. This effect occurs from saturation of water protons associated with macromolecules, particularly those hydrogen-bonded to macromolecules. The degree of magnetization transfer is greater with fast spin echo sequences than conventional spin echo sequences although it occurs to some degree in both [8]. It may be more acute due to the very rapid deployment of 180° pulses in HASTE type sequences.
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Fig. 1. FSE sequences showing basic tissue contrast in the pelvis. (A) T2-weighted conventional spin echo image (TR/TE= 2350/90) in the sagittal plane through the midline shows the normal zonal anatomy of the uretus with, hyperintense endometrium, low signal intensity junctional zone, and higher signal intensity outer myometrium. (B) T2-weighted sagittal FSE sequence (TR/TEeff =3000/90) shows the same basic tissue contrast as in (A) but was performed in 3 min and 24 s. (C) T2-weighted sagittal single shot FSE sequence (TR/TEeff =25338/98) was performed in 25 s for an equivalent number of slices. Note that the signal intensity basic tissue contrast is the same as in (B) but there is slight blurring of the soft tissues of the pelvis. However, note the sharp delineation of the fluid in the bladder, vaginal fornix and endocervical canal (arrowheads).
The development of high performance gradient systems has improved fast spin echo sequences. Deployment of faster gradients has enabled shorter echo spacing, that is the time between the 180° pulses, and thus the time between the signals. This has several beneficial effects. Firstly, reducing the echo spacing reduces the time allowed for T2 decay from phase encoding step to phase encoding step. This effectively reduces image blurring and shortens the overall time it takes to deploy the echo train. Decreasing the echo spacing also has the effect of diminishing motion artifacts [9]. Lastly, improved gradient systems make acquisition of longer echo trains, such as those used in single shot techniques described in the next section, more practical.
2.2. Single shot fast spin echo HASTE images were the first type of sequence to obtain high quality T2-weighted spin echo-type images in a single breathhold. Previous attempts to obtain T2-weighted images in a breathhold included T2-weighted gradient echo images such as steady state free precession. These suffered from low resolution and considerable magnetic susceptibility artifacts that limited their use in the abdomen and pelvis. HASTE sequences employ half-Fourier acquisition and a long echo train fast spin echo sequence to obtain all the data necessary to form k-space from a single acquisition or a single ‘shot’. Therefore, these sequences are referred to as HASTE or single shot fast spin echo.
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Fig. 2. Fast spin echo compared to single shot fast spin echo in a woman with uterine myomas. (A) Fast spin echo image sequence (TR/TEeff = 3233/90) through the uterus shows muliple uterine myomas of varying signal intensity. (B) Single shot FSE sequence (TR/TEeff = 17137/98) shows lower signal-to-noise (SNR) overall. There is blurring of the tissues of the uterus and myomas as well, but the urine in the bladder remains sharp.
Single shot fast spin echo sequences can be used to obtain relatively motion free breathold sequences in the pelvis with tissue contrast similar to more conventional fast spin echo. For conventional pelvic imaging, the advantage of these sequences is the ability to image the abdominal wall, mesenteric fat and bowel without motion. The main disadvantage of these sequences is the overall lower SNR and the tissue blurring in the phaseencoding direction due to alterations to the pointspread function, as previously described for fast spin echo sequences. These effects are more pronounced on single shot fast spin echo sequences and are the major factor limiting their more widespread use in the abdomen and pelvis (Fig. 2). Single shot fast spin echo sequences have been substituted for fast spin echo sequences for imaging the pelvis (Fig. 2). For certain specific, ‘simple’ applications such as identification of fibroid, single shot fast spin echo sequences may be able to obtain a study in a shorter period of time [10]. Imaging of the bowel is superior with single shot fast spin echo sequences, for two reasons [3]. Imaging time is sufficiently short so that perstalsis-related motion artifacts are reduced. Also, the long T2 of the luminal contents will cause the mucosal surfaces to be imaged with high resolution.
2.3. Echo planar imaging Echo planar imaging has not been widely used for imaging the pelvis. The reasons for this are several fold. First, until recently, gradient performance has not been
up to the task of the rapid gradient switching that is necessary to generate echo planar images on standard MR imaging systems. Secondly, magnetic susceptibility and chemical shift artifacts are a problem with echoplanar imaging and would greatly limit its potential in the pelvis. Magnetic susceptibility is high at interfaces between gas and soft tissues. This situation occurs commonly in the tortuous sigmoid colon, small bowel, and rectum and therefore, structures adjacent to these may be distorted or obscured on echo planar images. Lastly, it has not been considered necessary to obtain the degree of temporal resolution that echo planar imaging offers. In general, for imaging the pelvis, spatial resolution is at a premium. Contrast resolution or tissue contrast is less important and temporal resolution may be the least important of all. Some degree of temporal resolution is important when studying dynamics of contrast enhancement and vascularity, but this type of information has rather restricted utility in the pelvis.
2.4. Obstetric imaging MR images of the intrauterine fetus have always been susceptible to motion induced artifact [11,12]. This motion has two sources. One is respiratory induced movement of the uterus due to maternal or respiratory motion. The second is gross movements of the fetus itself. Motion of the fetus causes motion of the amniotic fluid around it, which generates a marked ghosting artifact and blurring such that, in order to image the fetus for possible fetal abnormalities, paralysis of the
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Fig. 3. Single shot FSE imaging of gravid uterus. (A) Single shot FSE sequence (TR/TE 19062/98) shows excellent anatomic detail in the fetus. Note the identification of the liver (li) lungs (lu) and the severe hydraencephaly (h). The low signal intensity vessels of the umbilical cord lie on either side of the fetus’s neck. (B) Axial single shot FSE sequence with the same parameters as (A) shows hydraencephaly with some preservation in a vascular territory of the occipital lobe (arrow). Note the delineation of the placenta posteriorly (p).
fetus was frequently necessary with conventional spin echo sequences or even fast spin echo sequences. Fast spin echo sequences are suitable for imaging the placenta and uterus but frequently give poor quality images of the fetus itself due to motion induced artifact. The single shot sequences (HASTE) have dramatically improved the potential for MR imaging of the obstetric patient. These sequences are obtained in a breathhold but this is not nearly as important as the overall speed of imaging. What is more critical is that these images be acquired quickly before gross fetal motion has occurs. Frequently some of the images will be degraded if motion occurs during their acquisition; whereas others will be relatively free of motion artifact when acquired in the absence of fetal motion. Therefore, repeated acquisition of the sequences can be performed to obtain some diagnostic images that are free of motion artifact. The blurring of soft tissue structures demonstrated in HASTE sequences is of lesser importance when imaging the fetus because most structures are bordered by fluid, and therefore image sharply. Examining the fetus with HASTE imaging has been made easier by obtaining oblique planes that represent anatomic cross sections realtive to the fetus, not the mother (Fig. 3). Single shot fast spin echo sequences are useful for imaging the fetus or placenta in cases where ultrasound findings are equivocal. The display of fetal anatomy before 20 weeks of gestational age is usually inferior to that of ultrasound, due to motion and the small size of the fetus [13,14]. Single shot fast spin echo images show
particular advantage compared to ultrasound in situations where ultrasound is limited: in the setting of oligohydramnios and in late third trimester. MR imaging can demonstrate additional findings in fetuses with neurologic abnormalities, such as germinal matrix hemorrhages and agenesis of the corpus callosum [15–19]. MR imaging is also useful in showing the contents of diaphragmatic hernias and omphaloceles because liver can easily be distinguished from bowel [20]. Widespread use of the single shot fast spin echo sequence for fetal imaging awaits research into the optimal indications for MR imaging in this population. Another approach to imaging the fetus uses echo planar imaging [21]. While echo planar imaging can easily achieve images in less than a second like single shot fast spin echo, SNR considerations probably favor the latter. Echo planar imaging has been used to investigate aspects of lung and liver maturation [22,23].
2.5. MR urography T2-weighted fast spin images can be performed with a sufficiently long TE to effectively image only fluid [24]. This sequence is termed MR urography when applied to the pelvis and MR cholangiography when applied to the biliary tree. The basic principle is the same: to acquire an image or images which display the entire fluid-filled collecting system. This can be accomplished by one image representing a slab thick enough to include the renal pelvis, ureters and bladder [25,26]. An alternative is to acquire contiguous thin-section images and perform a computer generated reconstruc-
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Fig. 4. FSE sequences used for MR urography. (A) Axial T2-weighted FSE image through the base of the bladder (TR/TEeff =4350/119) shows a tumor invading through the muscularis of the bladder. (B) MR urogram made from 3 mm thick coronal sections using a FSE sequence (TR/TEeff = 12 000/260, 4 NEX) shows bilateral ureters. The right ureter is dilated down to the level of the tumor mass, but the site of the obstruction (arrow) is not clearly delineated on this anterior to posterior projection. Note the right hydronephrosis. (C) Cropped view of the distal right ureter shows the dilated ureter (arrowheads) terminating at the level of the tumor at the ureterovesicle junction (arrow). This is a lateral projection view of a urogram made sagittal images.
tion (e.g. maximum intensity projection or MIP) from multiple perspectives to visualize the urinary system (Fig. 4). The single section slab approach has the advantage of fast acquisition, which allows acquisition in suspended respiration [25,26]. MR urograms can show the site of urinary tract obstruction very well but because the unobstructed ureters are intermittently peristalsing, they are generally not shown in their entirety. Unfortunately image resolution is not equivalent to intravenous urograms or retro-
grade pyelograms, so they cannot be used as a substitute for these in many applications. In particular, the detection of transitional cell neoplasia and small calculi may not be optimal on these images.
3. Ultrafast vascular imaging in the pelvis Vascular imaging of the pelvis has always been a major challenge for body vascular imaging using conventional
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Fig. 5. A 71-year-old man with femoral–femoral cross graft. Time-of-flight (TOF) imaging compared to 3D gadolinium enhanced imaging in the pelvis. (A) Anterior to posterior projection reconstruction of individual axial 2D TOF gradient echo sequences (TR/TE 33/6.4), flip angle of 60° shows the graft entering the right common femoral artery and bilateral occulsion of the native iliac arteries. The femoral – femoral cross graft is not well shown due to inplane flow saturation. (B) 3D gadolinium enhanced gradient echo acquisition (TR/TE 11.2/2.1) flip angle of 60° shows greater anatomical detail in the arteries compared to (A). Note that the femoral – femoral cross graft (arrowheads) is well shown with the 3D gadolinium enhanced technique. Also note that the left external iliac artery (arrow) is shown in (B) but is not shown in (A). This is due to reversal of flow with filling via the cross graft.
techniques. Of the various MRA sequences available, time-of-flight (TOF) gradient echo imaging was used most widely prior to the development of faster imaging methods. TOF imaging generally works well for imaging the pelvic veins, whereas the pelvic arteries have provided a greater challenge [27]. This is partly due to the fact that prominent arterial pulsations are normally present in the pelvic arteries, leading to marked differences in flow-related enhancement from phase encoding step to phase encoding step. The changes in flow-related enhancement leads to prominent ghosting artifacts and a resultant loss of signal intensity in the pelvic arteries. Additionally, the pelvic arteries are frequently tortuous and thus susceptible to in plane flow saturation on TOF images (Fig. 5). Lastly, the pelvic arteries were vulnerable to magnetic susceptibility artifact resulting from neighboring colon. Because the pelvic arteries are of critical imaging importance in evaluating patients with claudication or tissue loss, precise delineation of any stenoses is crucial. For these reasons, optimal imaging of pelvic arteries was not possible before the development of fast imaging techniques with gadolinium enhancement [27]. 3D MRA with gadolinium proved to be a major advance in overcoming the special problems posed by pelvic arteries [28]. With the short echo times used in 3D MRA, magnetic susceptibility artifacts were no longer a significant source of error. The dramatically shortened T1 of blood provided by gadolinium eliminates the in plane flow saturation that was frequently
seen in torturous arteries (Fig. 5). Lastly, because the 3D MRA technique relies on imaging blood by virtue of its shortened T1 and not flow-related enhancement per se, this technique is not susceptible to the alterations in flow-related enhancement which led to the view-to-view errors and pulsation artifacts of the 2D TOF technique [28]. The details of the gadolinium enhanced 3D gradient echo imaging technique have been described elsewhere [28]. Briefly, this technique involves a 3D gradient echo acquisition to image a volume with thin sections amenable to 3D reconstruction and reformatting. This type of sequence can be performed on more conventional gradients systems; however, the high performance gradients with higher gradient strengths and slew rates allow the 3D sequence to be performed in a shorter time period. Acquisition of the 3D sequence during a shorter time period means that timing the acquisition to the bolus is less problematic, and the sequence can be performed with breathholding to eliminate some sources of motion. For imaging a given volume, an orientation to the slab must be chosen. This differs from TOF MRA, where the pelvis is imaged in cross-section (axially) and final reconstructions are made in the anterior–posterior projection. Thus, with TOF MRA final resolution depends on slice thickness. For imaging the pelvis with a 3D volume acquisition, the coronal plane or slab includes the target vessels, and the section thickness determines the final resolution. The 3D gradient echo
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Fig. 6. Acquisition of more than one phase in the 3D gadolinium enhanced MR angiography sequence. A 73-year-old woman with infected femoral graft and venous thrombosis. (a) 3D gadolinium enhanced projection reconstruction of a coronal slab from a gradient echo sequence (TR/TE 13/2.3) shows a completely occluded iliac arterial system. The right arterial system is severely diseased with multiple stenosis. A cross graft femoral to popliteal graft is seen (arrowheads). (B) The venous phase of the 3D gadolinium enhanced technique shows enhancement around the graft focally (arrowheads) which prove to represent infection in the graft. Note the numerous venous collaterals from bilateral iliac deep venous thrombosis.
sequence is obtained with minimum TE and TR. Because the large number of phase encoded steps involved in the acquisition, the minimum TR obtainable will dramatically effect the total acquisition time. Gadolinium is injected intravenously and the plateau of arterial phase enhancement is timed to correspond to the center of k-space acquisition of the 3D gradient echo scan. If the 3D sequence is a half-Fourier acquisition, then the center of a k-space may occur at the beginning of the acquisition rather than in the middle. 3D gadolinium enhanced sequences are less susceptible to artifacts in general than 2D TOF imaging. Several studies have shown greater accuracy with the 3D gadolinium enhanced technique compared to the 2D TOF technique [29 – 33]. Pitfalls in the 3D gadolinium enhanced technique are primarily related to technical problems with the performance of the examination. Because gadolinium dramatically shortens the T1 of the blood, selective in-flow saturation of venous vessels is far less effective with the gadolinium enhanced technique compared to the TOF technique. This is compounded by the fact that the usual acquisition for the gadolinium enhanced sequence is in the coronal plane. Therefore, saturation bands for suppression of venous in-flow is not possible, and selective imaging of the arterial systems depends upon critical timing so that venous enhancement does not occur. A corollary to this is if the sequence is rapidly repeated after the arterial
phase then a venous phase ‘venogram’ may be obtained (Fig. 6). 3D gadolinium enhanced MRA is particularly well suited to patients with complex arterial grafts in place. Imaging of femoral bypass grafts was a particular problem with the 2D TOF imaging due to in-plane flow saturation. With the 3D gadolinium enhanced technique, the flow direction is irrelevant and therefore, flow in grafts running in any direction can be effectively imaged (Fig. 5). Surgical clips may still cause artifactual stenosis with the 3D gadolinium enhanced sequence, but the problem will be less profound compared with 2D TOF imaging because the TE of the 3D sequence is so much shorter. 3D gadolinium enhanced sequences have been used in a number of applications in the pelvis [33,34]. They are more accurate for identification of stenoses in the iliac arteries than conventional TOF sequences. They are useful for identification of stenosis and thrombosis affecting renal transplants, arteries and veins [35]. Because of the short transient time of arterial blood through the kidneys, it is relatively simple to image both the arteries and veins with this type of sequence. These sequences can detect arterial lesions associated with impotence [36]. The potential of these sequences for delineation of arterial supply in other pelvic disorders per se has not been explored.
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Fig. 7. Dynamic gadolinium enhancement of the uterus in a woman with cervical carcinoma. Sagittal T2-weighted fast spin echo image (TR/TE= 3883/126) shows infiltration of the cervix (arrow) with slightly higher than normal tissue signal intensity. (A) Dynamic gradient echo multiplanar images (TR/TE= 125/2.5) performed before (B), in the arterial phase (C), and venous phase of gadopentatate dimeglumine enhancement. The cervix, which normally enhances less than uterine corpus, is enhancing more than corpus myometrium, indicating diffuse infiltration with tumor.
4. Fast gadolinium-enhanced imaging in the pelvis Dynamic imaging with contrast enhancement is a valuable tool in several areas of the body, most prominently breast, liver and pancreas. Dynamic imaging of the pelvis has been explored to some extent and is primarily useful in the differentiation of scar versus recurrent tumor in the postoperative pelvis and for evaluation of endometrial abnormalities and bladder tumors (Fig. 7) [37–44]. It is probably not useful in evaluation of prostatic disorders or adnexal disorders, although it has not been thoroughly explored for these applications.
For the purpose of dynamic images, three parameters are frequently in conflict with one another in providing the maximal information possible. These are: temporal resolution (i.e. the number of images of a given area obtained per unit time), spatial resolution, and slice coverage (or the volume of tissue that is scanned). The best temporal resolution from low-resolution images are obtained at one slice position, while repeating the image rapidly during contrast administration. This sequence has been used for imaging prostatic lesions, individual lymph nodes, and cervical and bladder tumors (Fig. 7). The problem with this approach is that it gives information about only one plane rather than the
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Fig. 8. Three-dimensional gadolinium-enhanced imaging of the pelvis in a 54-year-old woman with a cystic mass of the left ovary. (A) Sagittal T2-weighted fast spin echo image (TR/TEeff = 3250/105) shows a multilocular cyst of the left ovary (arrow). (B) Sagittal fat-suppressed 3D gradient echo image (TR/TE=7.0/2.3 ms) is performed with 32 4 mm thick sections and a 15° flip angle in 26 s. Note the sharp delineation of the bowel (white arrow). (C) Same sequence after gadopentatate dimeglumine administration shows enhancement of the cyst walls (arrow) and pelvic vessels. (D) For comparison, a 2D gradient echo image (TR/TE= 400/2.1) yields higher signal-to-noise (SNR) and resolution but is not performed with suspended respiration. Local sigmoid wall thickening from Crohn’s disease is fistulized (thin arrow) to the left ovary (black arrow).
pelvis or tumor as a whole. Another approach to dynamic imaging involves using higher resolution images with a longer TR to allow multiplanar imaging of much if not all of the pelvis. However, the temporal resolution of this type of imaging is much lower. Generally, images corresponding to arterial phase, venous phase, and parenchymal or interstitial phases are obtained. With faster and more powerful gradients, dynamic imaging with 3D T1-weighted acquisition becomes possible. This type of sequence can be used for dynamic
gadolinium enhancement of organs in the arterial, venous, and parenchymal phases. There are two reasons for performing 3D acquisitions for dynamic scanning. Firstly, with similar parameters, 3D acquisitions yield higher SNR acquisitions. Partly for this reason, thinner sections can be obtained without cross talk, providing higher spatial resolution in the slice select direction (Fig. 8). Secondly, 3D acquisitions can be reconstructed to provide MR angiograms in the arterial phase since these sequences are fundamentally similar to the 3D gadolinium enhanced MR angiography sequences.
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There are some disadvantages to using the 3D gadolinium enhanced technique for dynamic imaging of the pelvis: (1) there is competition between the parameters necessary for high resolution imaging of the pelvis, e.g. a 20 cm field-of-view may not be implemented with dynamic images due to unacceptable demands on the gradients necessary to obtain the proper TE. Secondly, the 3D sequences use a very short TE and TR. The short TR lends very little T1-weighted contrast to the image (Fig. 8). Thus, relative advantages and disadvantages must be considered for a given clinical application before showing 3D gadolinium enhances images over 2D gradient echo images.
5. New directions The potential of newer techniques, such as perfusion or diffusion imaging in MR imaging of the pelvis, has been largely unexplored. In the abdomen, measuring the apparent diffusion coefficient (ADC) may help discriminate hemanigomas from solid tumors [45 – 47]. Chronic renal failure may be associated with alterations in the ADC as well [48]. However, little is known about the state of water diffusion in pelvic organs. Yang et al. found that in rats water diffusion is anisotropic in the myometrium, but not in the endometrium [49]. Apparent water diffusion coefficients depend on hormonal influences in the endometrium, but not in the myometrium. These preliminary animal studies suggest that diffusion imaging of the human uterus may be useful in studying endometrial physiology [49]. Fast spin echo sequences can be modified with diffusion-sensitizing gradients [50]. These sequences can also be used to study perfusion with arterial spin labeling [51]. Practical and convenient fast sequences such as these may spur research into the practical applications of diffusion and perfusion imaging in the pelvis. Ultrafast imaging is important in real-time MR ‘flouroscopic’ techniques that may prove useful for interventional procedures [52,53]. The clinical potential of MR interventional procedures in the pelvis has not been explored. Possible applications include interventional thermal ablation of tumors, biopsies, and radiation seed placement [54].
6. Summary Ultrafast imaging is made possible by stronger gradients capable of generating higher slew rates and thus much more rapid deployment of gradients. These gradients on conventional MR systems (Siemens Vision, GE EchoSpeed, Picker EDGE, Philips ACSNT) accelerate and improve pulse sequences that are important in imaging the pelvis, such as fast spin echo, fast spoiled
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gradient echo, and 3D fast gradient echo. Other sequences such as single shot fast spin echo (HASTE) and echo planar suffer from lower spatial resolution and certain artifacts which limit their effectiveness in the pelvis, where spatial resolution is critical. These sequences could be useful for diffusion imaging, the potential of which is largely unexplored in the abdomen. 3D gadolinium-enhanced MRA techniques are generaly superior for imaging the pelvic arteries.
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