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Magnetic Resonance Imaging in Urology
0022-534 7 /84/1324-0641$02.00/0 Vol.
TBE JOURNAL 8F UROLOGY
Copyright© 1984 by The 'Williams & Wilkins Co.
Printed
October
U.S.A.
Review Article MAGNETIC RESONANCE IMAGING IN UROLOGY RICHARD D. WILLIAMS*
AND
HEDVIG HRICAK
From the Departments of Urology and Radiology, University of California School of Medicine and Veterans Administration Medical Center, San Francisco, California
Although still m its developmental stages, magnetic resonance imaging is expected to become one of the most significant advances in diagnostic body imaging. This optimistic view is based on the unique ability of magnetic resonance imaging not only to provide images of precise anatomical detail in virtually any planar projection desired but also to deliver simultaneously biochemical information concerning the tissues examined. Importantly, all information is obtained without imparting any known harmful effect on the subject, that is no ionizing radiation is used and iodinated contrast medium enhancement is unnecessary. The purpose of this review is to familiarize the urological surgeon with the underlying principles of magnetic resonance imaging and to present the current state of the art with respect to anatomical areas of interest to the urologist. MAGNETIC RESONANCE IMAGING-BASIC PHYSICAL PRINCIPLES
Even though the physical principles of magnetic resonance were described first in 19461 · 2 and the concept of magnetic resonance body imaging for medical use existed in the early 1970s3 - 8 success in adapting magnetic resonance for clinical use awaited development of the computers and software required for reconstruction of the i;;;l0 8 data bits necessary to produce x-ray computerized tomographic (CT) imaging. The fundamental concepts of magnetic resonance have been reported in detail recently and are based on the presence of atomic nuclei with an odd atomic number in living tissues. 8- 16 Examples of such atoms include 'hydrogen (1H), 13 carbon, 17 oxygen, 19fluorine (1 9F), 23 sodium (2 3 Na) and 31 phosphorus (31 P). Of these, hydrogen atoms (protons) are the most abundant based on the ubiquity of water (which contains 2 protons) and lipids within living tissues. Because these atoms have an electric charge and an intrinsic property of spin or angular momentum, each pro· duces a small magnetic field around it that can be described a vector (fig. 1, A). In undisturbed tissue the net magnetic vector of all protons present is zero (fig. 1, owing to the random orientation of the individual magnetic vectors. However, if a group of protons is placed in a strong external magnetic field, the magnetic vectors of the protons will tend to align in that field (fig. 1, C). Generally, they orient either parallel or anti-parallel to the external magnetic field. More protons will line up parallel (because this represents a lower energy state) and their net magnetic vector will be in the direction of the north pole of the external magnetic field. Magnetic field strengths are referred to in either Gauss or Tesla units (LO Tesla unit = 10,000 Gauss units). For example, the gravitational field of earth is approximately 0.6 Gauss units, whereas the field strengths used for magnetic resonance imaging are much stronger (0.1 to 1.5 Tesla units). Following alignment within the magnet each still-spinning proton begins to spin (precess) about the axis of the magnetic field, similar to a *Requests for reprints: Urology, U-518, University of California, San Francisco, California 94143.
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spinning top rotating around its own axis while rotating simultaneously around the vertical axis of the gravitational field of earth. The frequency of precession is characteristic for the external magnetic field strength. If energy in the radiofrequency range is applied at a right angle to the axis of the magnet at exactly the frequency of the protons' precession, their individual vectors merge because the added energy causes a rapid, repetitive unified flipping of the protons between the parallel and anti-parallel orientation (fig. 1, D). This change is termed resonance. The net magnetic vector of the protons will be displaced to a new position by the radiofrequency pulse. Because the protons absorb radiofrequency energy during this process, when the radiofrequency pulse ceases the absorbed radiofrequency energy is released as the protons' magnetic vectors return to their original state (fig. 1, E). The radiofrequency energy released will be emitted at exactly the same frequency as that of proton precession and can be detected by a radiofrequency wave monitor. It is this measurable release of energy that is called the magnetic resonance signal. The intensity of the magnetic resonance signal is proportional to the relative number of the atomic nuclei resonating. In living tissues only the atoms present in relative abundance can be used to provide a magnetic resonance image. 'H, 19F, 23 Na and 31 P always are present in these isotopic forms in living tissues, yet only 1 H is present in enough concentration and emits a strong enough signal for current use in clinical magnetic resonance imaging. Therefore, to date only proton magnetic resonance imaging has medical relevance. For medical magnetic resonance imaging we are more interested in the relative position or spatial relationship of 1 H atoms within the tissue being examined than merely their presence and relative numbers. To provide this information it is necessary to vary the strength of the external magnetic field, which is accomplished by 1m,rouu1.:1r1g another magnetic force within the aperture of the external. magnet so that a field gradient can be produced. Because the precessional frequency of each proton is related directly to its position in the graded external magnetic field the radiofrequency signal emitted by the protons following excitation contains spatial information. Therefore, by application of radiofrequency pulses of different duration and different frequency, the resultant magnetic resonance signals produced also will be slightly different and, when transformed and reconstructed by computer, can provide an image depicting the relative position, density and biochemical environment of the protons. Therefore, this information can reveal subtle differences within and between tissues. In fact, these differences can be measured more precisely than is possible in CT imaging and, therefore, can produce images with potentially greater anatomical discrimination. It is this variability in the field strength and, thus, excitation of selected protons at a given radiofrequency pulse within the tissue that allow image formation in almost any plane desired. Besides proton density images, at least 3 other magnetic resonance parameters can be used to impart different infor-
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A
B
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FIG. 1. A, hydrogen atom (proton) with atomic number of 1 results in angular momentum and magnetic vector. B, random distribution of protons in tissue with net magnetic moment of zero. C, proton magnetic vectors undergo alignment within strong external magnetic field. D, radiofrequency (RF) waves at right angles to magnetic field cause proton vector realignment (precession). E, discontinuation of radiofrequency waves permits proton vectors to return to original alignment in magnetic field, which releases radiofrequency energy (NMR, nuclear magnetic resonance signal), which then is processed by computer to produce image.
mation: Tl and T2 relaxation times, and bulk motion of hydrogen or flow. The first 2 parameters are based on measurements of proton precessional relaxation times. Tl refers to the time required for proton precession to return to thermal equilibrium as evidenced by the tendency for alignment (discussed previously) parallel to the external magnetic field after cessation of each radiofrequency pulse. This time can vary considerably: in solids at low temperatures, when random thermal motion is minimal, Tl may be long (hours), whereas in pure liquids Tl may be only several seconds. In liquids containing proteins (such as tissue fluids) Tl is even shorter. During image formation an interval (recovery time) must be introduced between successive radiofrequency pulses so that recovery of magnetization of protons after the radiofrequency pulse ceases can occur and be measured. Increasing recovery time can increase the intensity of the magnetic resonance signal but also can increase the scanning time. Therefore, recovery time must be selected carefully for medical imaging. T2 relaxation time refers to the signal decay time induced by the magnetic forces exerted on the protons in their microenvironment (affected by adjacent protons), which dampen coherent resonance until the signal ceases. T2 is measured in milliseconds and Tl in seconds. The interval between an applied radiofrequency pulse and signal reception is a measure of the role of signal decay and is designated by the instrument parameter (echo delay time). In solids T2 is in the microsecond range, while in liquids T2 is longer. Thus, variable instrument setting and the salient tissue characteristics affect the magnetic resonance signal intensity and images can be reconstructed, which relate to proton density, and Tl and T2 relaxation times. Tl and T2 images are grayscale maps synthesized from the different combinations of recovery time and echo delay time used to sample intensity of
the magnetic resonance signal at different points in time. Since approximately 50 msec. are needed for the proton signal to register, if the protons move faster through the plane of imaging, their signal will not be detected. This permits assessment of blood flow and represents the third magnetic resonance parameter, proton bulk motion. Currently, there are 4 major techniques used to obtain magnetic resonance intensity images: 1) partial saturation recovery, 2) free inductive decay, 3) spin echo and 4) inversion recovery. At our university spin echo and inversion recovery data acquisition techniques are used routinely. In the spin echo technique the image is dependent on the Tl and T2 relaxation parameters, hydrogen density (spin density) and blood flow. The instrument parameters that can be varied to produce tissue contrast with the spin echo technique are recovery time and echo delay time. Recovery time represents the interval between the repetitive radiofrequency pulses applied to the image volume and it can be varied from 0.5 to 2 seconds. Echo delay time indicates the interval from excitation of the nuclei by radiofrequency pulse and the receipt of the signal from the nuclei. Images are obtained at echo delays of 28 and 56 msec. in each section plane. In the inversion recovery technique the image is dependent on hydrogen density and Tl relaxation values. The instrument parameters that can be varied to produce tissue contrast are recovery time, inversion time and, at our institution, echo delay time as well. Recovery time is referred to as the interval between repetitive 90-degree radiofrequency pulses applied to image volume and it can be 1 or 1.8 seconds. Inversion time is the interval between initiation of 180 and 90-degree pulses, and it can be 420 or 277 msec. Echo delay time represents the interval between excitation of the nuclei by radiofrequency pulse and receipt of the signal from the nuclei. On inversion recovery, similar to spin echo, images are obtained at echo delay times of 28 and 56 msec. The necessary components of a magnetic resonance imager are a magnet capable of imposing a strong uniform magnetic field, gradient coils that can alter the magnetic field internally depending on the technique used, a transmitter to deliver radiofrequency waves to the subject, a receiver to gather magnetic resonance signals from the subject and deliver them to appropriate digitizing circuitry and, finally, a computer capable of processing the information received into a spatial display (fig. 2). Three different types of magnets are available: 1) permanent, 2) resistive and 3) superconducting. Currently, 2 major approaches provide magnets of sufficient strength for medical magnetic resonance imaging. The resistive coil electromagnet can produce useful images at a reasonable cost (approximately $800,000). However, the homogeneity of the field produced in this device (an important factor for image accuracy) cannot be greater than a few parts in 105 and, because the power requirements and thermal dissipation needs for this type of magnet are great, the maximal field strengths reachable range from 1 to 2 KGauss units (0.2 Tesla units). The superconductive magnet has the advantages of greater obtainable field strength (theoretically up to 14 Tesla units), superior homogeneity (1 part in 108 ) and increased stability. The basic magnet consists of wound filaments of a superconductor that are supercooled by liquid nitrogen and helium. Electricity is required to initiate the magnet but in the supercooled environment the superconductor has almost no resistance and, therefore, can conduct and maintain the current necessary to produce the magnetic field indefinitely. Therefore, the power requirements of the magnet are minimal. However, the initial cost of the superconductive magnet is approximately double that of the resistive device. Because superconductive magnets produce a field stronger and more stable than that of resistive magnets, and because possible imaging of nuclei other than protons will require a
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FIG. 2. Typical superconductive magnetic resonance imager. A, external view. B, longitudinal schematic view of components necessary for magnetic resonance image construction.
strong magnet in the future, the superconductive magnet at present has the preferential advantage. ADVANTAGES OF MAGNETIC RESONANCE IMAGING
The enthusiasm attending the emergence of magnetic resonance imaging for medical use is predicated upon certain advantages over current conventional radiological imaging techniques that already have been established: 1) magnetic resonance imaging avoids the use of ionizing radiation, 2) magnetization in the power range used is without harmful biological effects, 3) spatial resolution comparable to the latest generation x-ray CT imaging already is demonstrable, 4) contrast resolution is vastly superior, 5) iodinated contrast medium is not required, 6) no beam hardening or bone artifact (particularly important in imaging the pelvis) is seen, 7) blood flow measurement is possible, 8) imaging of 31 P, 19F or 23 Na may permit determination of cellular or tissue biochemical abnormalities and 9) image distortion owing to metal objects (such as surgical clips) is absent. The development of any new diagnostic modality also must be scrutinized with respect to possible risks, toxicity or disadvantages. The initial cost of the magnet, slow scanning time (compared to CT), and lack of detection of cortical bone and tissue calcium appear to be the only deterrents of magnetic resonance imaging. Similar to the early problems with CT, claustrophobia has been a problem in approximately 5 per cent of the patients examined. In vitro cellular toxicity studies have revealed no evidence of inhibition of deoxyribonucleic acid (DNA) metabolism and DNA function remains unaffected. 17• 18 Human studies have shown no short-term or medium-term effects despite multiple exposures of subjects and operator personnel. 19 The experience in our laboratory has shown no observed toxicity in > 1,000 patients with a wide variety of clinical conditions, thereby eliminating many of the limitations reported previously. 20 • 21 CLINICAL MAGNETIC RESONANCE IMAGING
The images presented were obtained on our magnetic resonance imager, a superconducting electromagnet functioning at
a field strength of 0.35 Tesla units (15 MHz.). 21 Images were produced via a selective irradiation technique with 2-dimensional Fourier transformation. The imaging sequences chosen in our laboratory were spin echo and inversion recovery. Spin echo is accomplished using radiofrequency repetition intervals of 500 to 2,000 msec., and echo delay times of 28 and 56 msec. These parameters allow as many as 20 anatomic sections to be obtained simultaneously in approximately 18 minutes. Inversion recovery is accomplished with repetition intervals of 1,000 and 1,800 msec., with an inversion to recovery of 420 and 280 msec., respectively, allowing up to 9 sections in 16 minutes. The choice of imaging sequences is important, since time limitations dictate that only 2 to 3 sequences can be accomplished at any one sitting to keep scanning time between 30 and 60 minutes. Because the art is evolving, the precise combinations of the variables required to produce discriminating images in each clinical situation have yet to be determined. MAGNETIC RESONANCE IMAGING GRAY SCALE
Magnetic resonance imaging is a function of mobile protons (spin density), Tl and T2 relaxation parameters, and bulk motion of protons (flow). Contrast on the magnetic resonance image is a relative phenomenon and is the result of inherent tissue characteristics, the imaging technique used (spin echo, free inductive decay, partial saturation recovery and inversion recovery) and the magnetic field strength used. Within each imaging technique tissue contrast is influenced by instrument parameters (recovery time, echo delay time and inversion recovery time). The gray scale in the spin echo magnetic resonance image shows fat as the brightest tissue followed by marrow and cancellous bone, brain and spinal cord, viscera, striated muscle, ligaments and tendons, rapidly flowing blood, compact bone and air in decreasing order of intensity. With the spin echo technique the signal acquisition, variations of echo delay time and recovery time will discriminate between tissue as a function of their relative length of relaxation times. The tissue contrast seen with the inversion recovery technique varies with the inversion recovery interval and recovery time. This technique
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FIG. 3. Normal kidneys. A, spin echo recovery time is 1 second and echo delay time is 28 msec. There is clear differentiation between higher intensity cortex ( C) and lower intensity medulla (M). Gerota's fascia (large arrow) is seen as line dividing perirenal from pararenal spaces. Small arrow denotes left renal vein. A, aorta. I, inferior vena cava. L, liver. B, inversion recovery. Differentiation between lower intensity medulla (M) and higher intensity cortex ( C) is more pronounced than on spin echo image. L, liver.
emphasizes variations in Tl relaxation time, while the spin echo technique is dependent on Tl and T2 relaxation times. NORMAL RENAL ANATOMY
The anatomic features of a normal kidney are defined clearly by magnetic resonance imaging (fig. 3). On spin echo images the renal hilus is imaged as a high intensity area owing to the presence of hilar adipose tissue. The renal veins and arteries are imaged as low intensity tubular structures traversing the hilar region. The low intensity of the vessel lumen is caused by rapid blood flow. The pelviocaliceal system and the ureters are imaged with low intensity owing to the long Tl and T2 relaxation values of urine. On images obtained during diuresis a mild dilatation of the pelviocaliceal structures can be appreciated. Within the renal parenchyma the renal cortex is distinguished easily from the medulla, since the cortex is of a higher intensity than the medulla yet lower than the perirenal fat (fig. 3, A). The medulla is a triangular low intensity area central to the cortex that has an intensity and prominence related to the state of hydration. 22 This fact suggests that imaging will be better with the patient in a hydrated rather than dehydrated state. The cortical medullary differentiation is enhanced on short recovery time pulse sequences and virtually disappears on images obtained with a long recovery time (2 seconds). Perirenal fat is seen as a high intensity area owing to the short Tl and long T2 relaxation times of adipose tissue. Gerota's fascia is visible readily as a low intensity line surrounding the perirenal space. On inversion recovery images the renal cortex is significantly brighter than the medulla and differentiation between the two is more evident than on spin echo images (fig. 3, B). Perirenal and hilar fat remains bright owing to a short Tl relaxation time. Urine remains dark because of the long Tl relaxation time (fig. 3, B). RENAL PATHOLOGICAL FINDINGS
FIG. 4. A, left ureteropelvic junction obstruction. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Right kidney is normal with excellent cortical medullary differentiation. Arrow shows normal right ureter. Calices ( C) of left kidney are dilated and there also is marked dilatation of left extrarenal pelvis (P). Cortical medullary differentiation of left kidney is not present. Note thickened Gerota's fascia (curved arrow). B, chronic glomerulonephritis. Both kidneys are small with marked thickening of renal parenchyma ( C). Simple renal cyst (arrow) arises from posterolateral aspect of right kidney. Note abundant hilar adipose tissue (asterisk).
Hydronephrosis is detectable easily by magnetic resonance imaging, as is a hydroureter. By varying spin echo instrument parameters the distended ureter can be differentiated from adjacent bowel, lymph nodes or blood vessels. On CT images contrast enhancement is required to make this differentiation. An example of hydronephrosis on magnetic resonance imaging is shown in figure 4, A. Because fat and fluid have opposite magnetic resonance imaging intensity signals, peripelvic cysts and sinus lipomatosis are distinguished readily by magnetic resonance imaging, whereas partial averaging on CT can cause ambiguity. In the diagnosis of renal parenchymal diseases magnetic resonance imaging promises to make a significant contribution (fig. 4, B). The increased sensitivity of magnetic resonance imaging to subtle changes in soft tissues suggests that monitoring the progression of medical renal disorders may be possible,
perhaps limiting the need for renal biopsy. In addition, because acquired renal cystic disease and renal neoplasms are increased in patients with chronic renal failure, magnetic resonance imaging, obviating the use of iodinated contrast agents, may prove efficacious. The ability of magnetic resonance imaging to distinguish solid from cystic renal lesions is accurate and comparable to that of contrast-enhanced CT. Cysts are of homogeneous low intensity and without internal structures (fig. 5). Tl values of cysts usually are prolonged and T2 values are >200 msec. A smooth interface with renal parenchyma usually is demonstrated. Cyst walls are not demonstrated. Hemorrhagic and infected cysts have been shown to have an increased spin echo intensity compared to simple cysts, suggesting that magnetic resonance imaging has a diagnostic advantage over other modalities in
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the evaluation of patients with renal trauma, or infected or hemorrhagic cysts (fig. Evaluation of renal tumors has shown a wide spectrum of magnetic resonance imaging findings (fig. 7, A and B). Spin echo images show a range from low to very high signal intensity lesions. 23 · 24 Prolongation of the Tl relaxation time of solid tumors relative to adjacent normal renal parenchyma is a common feature, whereas the T2 values are similar or increased relative to surrounding parenchyma. Staging of renal neoplasms by magnetic resonance imaging appears particularly promising in the evaluation of neoplastic involvement of the renal vein or inferior vena cava (fig. 7, C). A potential disadvantage of magnetic resonance imaging in the evaluation of renal masses is the lack of definition of calcification despite a variety of magnetic resonance imaging techniques. Calcifications have low intensity and frequently are not visualized. When seen, calcifications are not differentiated readily from dense fibrous tissue owing to a similar low intensity signal. Whether this feature of magnetic resonance imaging will prove to be a significant weakness requires further study. Although limited information is available to date it is expected that further experience will reveal magnetic resonance imaging to be of substantial value in defining renal parenchymal disease, detecting and categorizing renal masses, and staging renal neoplasms.
FIG. 5. Simple renal cyst. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Homogeneous low intensity cyst (C) is in right kidney. Cyst wall is not demonstrated but there is sharp demarcation at border with normal renal parenchyma (arrow). A, aorta. I, inferior vena cava. SMA, superior mesenteric artery. SMV, superior mesenteric vein. L, liver. S, spleen.
FIG. 7. A, renal cell carcinoma (T) in upper pole of right kidney is imaged with inhomogeneous high intensity. Low intensity line (arrow) interposed between tumor and renal parenchyma represents tumor pseudocapsule. L, liver. S, spleen. B, renal cell carcinoma is imaged as inhomogeneous predominantly low intensity lesion ( T) arising from medial aspect of right kidney. Arrow shows left renal vein. C, tumor thrombus (T) within inferior vena cava is imaged with medium intensity signal and is seen occupying entire lumen of inferior vena cava. Arrow shows left renal vein. Curved arrows show right and left renal arteries. There is hydronephrosis of right kidney. NORMAL ADRENAL ANATOMY
The adrenal gland is homogeneous and of low intensity on magnetic resonance imaging, and is surrounded and outlined clearly by the higher intensity periadrenal fat (fig. 8). The left adrenal is visualized routinely but the right adrenal is seen slightly less often. The left adrenal is an inverted V or a triangle, whereas the right gland appears superiorly as a thin line extending from the inferior vena cava juxtaposed between the ems of the diaphragm and the medial aspect of the right lobe of the liver, and inferiorly as a horizontal band posterior to the inferior vena cava. Differentiation of the adrenal cortex and medulla is possible, with the medulla being of lower intensity than the cortex. Subjectively, the adrenal medulla has longer Tl and slightly shorter T2 relaxation times than the outer rim. ADRENAL PATHOLOGICAL FINDINGS
FIG. 6. Polycystic kidney disease. Spin echo recovery time is 2 seconds and echo delay time is 28 msec. Both kidneys are enlarged markedly. Renal parenchyma is replaced by renal cysts of various size and intensity. Acutely hemorrhagic cyst is seen with high intensity (arrow).
Primary as well as metastatic adrenal lesions are detected easily by magnetic resonance imaging (fig. 9, A). Currently, differentiation between malignant and benign lesions is not possible. Since the size and volume of the gland can be dis-
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cerned adrenal hyperplasia is displayed easily by magnetic resonance imaging (fig. 9, B). These early results suggest that magnetic resonance imaging can match the ability of CT to demonstrate normal and pathological adrenal anatomy. Because of the ability of magnetic resonance imaging to differentiate adrenal cortex and medulla, and easy anatomical display of the adrenals (surrounding fat, and the adjacent liver and large blood vessels), magnetic resonance imaging eventually may surpass CT in the differential diagnostic evaluation of the adrenal gland.
nance imaging showed that the tumor was separate from the kidney and that the renal vein was patent. The leiomyosarcoma was inhomogeneous and contained a large high intensity central area documented at operation to be a liquefied hematoma. The proximal extent of the vena caval thrombus was capped by a high intensity lesion that proved to be a clot on the leading edge of the tumor. Collateral venous blood vessels coursing around the thrombosed vena cava were seen easily (fig. 11). These features were not appreciated even in retrospect by the CT studies.
RETRO PERITONEUM
Little experience has been gained to date in magnetic resonance imaging of the retroperitoneum other than that described previously in relation to renal and adrenal anatomy. 25 - 30 Retroperitoneal fibrosis is imaged as an irregularly shaped medium to low intensity mass. 31 The signal intensity is higher in the plaques than in adjacent muscle but less than that of adjacent fat. Tl values varied considerably but T2 values were low consistently. The extent of involvement is determined easily by magnetic resonance imaging and is striking on sagittal scans. Because of the lack of signal from vessels with flowing blood, the perivascular component and consequent vessel narrowing are demonstrated easily. Evaluation of the vascular involvement by retroperitoneal fibrosis is superior by magnetic resonance imaging than contrast-enhanced CT (fig. 10). We also have examined 1 case of a retroperitoneal leiomyosarcoma occupying the entire right retroperitoneal space with an occlusive tumor thrombus in the inferior vena cava. Magnetic reso-
FIG. 8. Normal adrenal gland. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Lateral and medial limbs of left adrenal gland (arrow) are imaged clearly. Within kidneys there is superb differentiation between lower intensity medulla (M) and higher intensity cortex (C). St, food-filled stomach. L, liver.
FIG. 9. A, adrenal metastases from carcinoid tumor (inversion recovery image) are imaged as low intensity lesions (arrow). Large metastatic focus (small arrows) is within liver (L). St, stomach. S, spleen. B, adrenal hyperplasia. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Both adrenal glands are enlarged (arrows).
FIG. 10. Retroperitoneal fibrosis. A, magnetic resonance imaging spin echo image recovery time is 1 second and echo delay time is 28 msec. Retroperitoneal plaque (P) encircles aorta. Normal flow through aorta (arrow) is demonstrated. Curved arrow shows inferior vena cava. Arrowhead shows stent within left ureter. B, CT scan shows retroperitoneal plaque encircling great vessels (P). Calcifications are within aortic wall. No information regarding flow is obtained. Inferior vena cava cannot be separated from plaque. Arrowhead shows stent within ureter.
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and urine is determined easily, and wall thickness and irregularity are appreciated readily. 33 Pelvic bones are distinguished easily by magnetic resonance imaging. Bone marrow is visualized with a high intensity signal. The surrounding cortical bone emits no signal owing to lack of mobile protons. If the bladder is not distended bowel loops can present confusing artifacts at the dome of the bladder. However, analysis of several anatomic sections with a variety of pulse sequence intervals will permit the proper distinction to be made. BLADDER PATHOLOGICAL FINDINGS
FIG. IL Leiomyosarcoma of inferior vena cava. Large right retroperitoneal mass has inhomogeneous intensity. High intensity area (H) represents liquefied hematoma. Tumor (T) invades entire vena cava.
Hypertrophied bladder wall has a medium intensity signal (fig. 13). Bladder tumors have a signal intensity higher than bladder muscle and, thus, can be distinguished easily from the normal bladder wall (fig. 14, A and B). Depth of penetration of bladder neoplasms is appreciated best by combining axial and sagittal images. Magnetic resonance imaging appears to show particular promise with respect to local staging. Also, magnetic resonance imaging is superior to CT in the evaluation of tumors located at the bladder base. NORMAL PROSTATE ANATOMY
The normal prostate reveals a homogeneous medium intensity signal on magnetic resonance imaging (fig. 14, C). With multiplanar images, contiguous structures, such as bladder base, rectum and periprostatic fat, readily are apparent. The seminal vesicles have a medium intensity signal and are clearly distinguished embedded within the pelvic fat (fig. 15). PROSTATE PATHOLOGICAL FINDINGS
Benign prostatic hyperplasia is imaged with a homogeneous medium intensity signal (fig. 16, A). Prostatic carcinoma appears to have a higher intensity signal than normal or adenomatous tissue (fig. 16, B). Involvement of the seminal vesicles (stage C), confirmed by palpation and pathological examination, is evident on magnetic resonance imaging. Distortion of the prostatic contour and disruption of the lower intensity surgical capsule are seen in patients with stage C disease. If these findings are confirmed when larger numbers of patients are studied magnetic resonance imaging will emerge as the modality of choice in the diagnosis and local staging ofprostatic neoplasms. In addition to the anatomic areas described, there is a superb display of penile and inguinal canal anatomy, with differential
FIG. 12. Sagittal scan of male pelvis shows Denonvilliers' fascia (arrow) separating prostate gland (P) from rectum (R). B, bladder. SV, seminal vesicles. CS, corpora spongiosa. CC, corpora cavernosa. SP, symphysis pubis.
PELVIS
Lack of respiratory motion, no bone or beam hardening artifact and superior contrast resolution make magnetic resonance imaging the superior modality in the evaluation of the pelvis. 32 NORMAL VESICAL ANATOMY
The bladder is distended routinely with urine during scanning and, therefore, is demonstrated easily (fig. 12). The signal intensity of urine is low. 32 Differentiation between vesical wall
FIG. 13. Bladder wall hypertrophy. Bladder wall (arrows) is thickened markedly and of homogeneous medium intensity signal that increases with longer recovery times. A, spin echo recovery time is 1 second and echo delay time is 28 msec. B, recovery time is 2 seconds and echo delay time is 28 msec. R, rectum.
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FIG. 15. Seminal vesicles (SV) are imaged with medium intensity signal posterior to bladder (B). R, rectum.
FIG. 14. A, transitional cell carcinoma ( T) is seen along right lateral and posterior walls of bladder (small arrows). Bulk of tumor protrudes within bladder lumen. Tumor also involves dilated right ureter (curved arrow). R, rectum. B, adenocarcinoma of urachal cyst (T) is imaged with medium intensity at dome of bladder. High intensity lesion outlining bladder lumen represents transitional cell carcinoma (asterisks). Foley catheter (arrow) is seen within urethra. P, prostate gland. SV, seminal vesicles. R, rectum. C, axial scan shows normal prostate (P) imaged with medium intensity signal. Small arrows show levator ani muscle. R, rectum.
demonstration of the cord structures, and excellent images of the testes. Pathological conditions of these areas have yet to be imaged. Magnetic resonance imaging of the pelvis appears to have contrast resolution capabilities exceeding that of CT, even with the use of contrast enhancement. Staging of pelvic neoplasms appears to be improved by the ability to examine the pelvis in multiple planes. Currently, the substantial obstacle appears to be the inability to depict pelvic lymph nodes clearly. However, this problem is expected to be resolved with experience and improvement in scanning techniques. FUTURE PERSPECTIVES
Our view of the potential of magnetic resonance imaging is optimistic. To date, the science is in its infancy. Nevertheless, for many organ sites the images produced by early generation magnetic resonance imagers already are equal in anatomical
resolution to those produced by current generation CT scanners. Furthermore, magnetic resonance images are superior to enhanced CT in other organs, most notably the brain, retroperitoneum and pelvis. With the advent of improved pulse sequencing techniques, perhaps specific for organs or disease states, the art should advance rapidly. In the near future it should be possible to obtain blood and urine flow measurements quantitatively, determine presence and volume of muscle ischemia or infarction and extend magnetic resonance imaging to other atoms, such as 31 P or 19F, which may allow specific physiological information to be gained. Perhaps a more exciting recent development in magnetic resonance imaging is the use of contrast agents to enhance the anatomic information now available. These agents are paramagnetic and, thus, shorten the Tl and T2 relaxation times. 34- 36 Manganese chloride already has been used to reduce magnetic resonance signal intensity from zones of myocardial infarction in isolated perfused canine hearts, 35 oral ferric chloride has been used successfully as an intraluminal bowel contrast agent, and inhaled 100 per cent oxygen has been used as an intracardiac blood magnetic resonance signal enhancer in human volunteers. Recently, nitroxide stable free radicals have been used effectively as an intravenous urographic contrast agent in animals. 37 If these developments are applicable routinely in humans, magnetic resonance imaging may well surpass currently available diagnostic modalities for selected organs and disease states. The ultimate role of magnetic resonance imaging in diagnostic medicine will require considerable research for definite conclusions to be drawn. However, in urology magnetic resonance imaging already has begun to establish its superiority in distinct anatomical sites and disease states. REFERENCES 1. Bloch, F.: Nuclear induction. Phys. Rev., 70: 460, 1946. 2. Purcell, E. M., Torrey, H. C. and Pound, R. V.: Resonance absorption by nuclear magnetic moments in a solid. Phys. Rev., 69: 37, 1946. 3. Jackson, J. A. and Langham, W. H.: Whole-body NMR spectrometer. Rev. Sci. lnstrum., 39: 510, 1968. 4. Damadian, R.: Tumor detection by nuclear magnetic resonance. Science, 171: 1151, 1971. 5. Lauterbur, P. C.: Image formation by induced local interactions: examples employing nuclear magnetic resonance. Nature, 242: 190, 1973. 6. Lauterbur, P. C.: Magnetic resonance zeugmatography. Pure Appl. Chem., 40: 149, 1974. 7. Hinshaw, W. S., Bottomley, P.A. and Holland, G. N.: Radiographic thin section image of the human wrist by nuclear magnetic resonance. Nature, 270: 722, 1977. 8. Pykett, I. L., Newhouse, J. H., Buonanno, F. S., Brady, T. J., Goldman, M. R., Kistler, J.P. and Pohost, G. M.: Principles of nuclear magnetic resonance imaging. Radiology, 143: 157, 1982. 9. Pykett, I. L.: NMR imaging in medicine. Sci. Amer., 246: 78, 1982. 10. Oldendorf, W. H.: NMR imaging: its potential clinical impact.
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FIG. 16. A, benign prostatic hyperplasia. Enlarged prostate gland (P) is imaged with homogeneous medium signal intensity. Foley catheter (arrow) is seen in place. B, bladder. SV, seminal vesicle. R, rectum. B, prostatic carcinoma. Enlarged prostate gland is imaged with inhomogeneous signal intensity. Area of high signal intensity anteriorly represents neoplastic tissue (arrows). Hosp. Pract., 17: 114, September 1982. 11. Brownell, G. L., Budinger, T. F., Lauterbur, P. C. and McGeer, P. L.: Positron tomography and nuclear magnetic resonance imaging. Science, 215: 619, 1982. 12. James, A. E., Jr., Partain, C. L., Holland, G. N., Gore, J. C., Rollo, F. D., Harms, S. E. and Price, R.R.: Nuclear magnetic resonance imaging: the current state. Amer. J. Roentgen., 138: 201, 1982. 13. Bradley, W. G. and Tosteson, H.: Basic physics of NMR. In: Nuclear Magnetic Resonance Imaging in Medicine. Edited by L. Kaufman, A. R. Margulis and L. E. Crooks. New York: IgakuShoin Medical Publishers, Inc., pp. 11-29, 1981. 14. NMR-An introduction. General Electric Company Technical Manual, 1981. 15. Loeffler, W. and Oppelt, A.: Physical principles of NMR tomography. Eur. J. Rad., 1: 338, 1981. 16. Partain, C. L., Price, R. R., Rollo, F. D. and James, A. E.: Nuclear Magnetic Resonance NMR Imaging. Philadelphia: W. B. Saunders Co., 1983. 17. Wolff, S., Crooks, L. E., Brown, P., Howard, R. and Painter, R. B.: Tests for DNA and chromosomal damage induced by nuclear magnetic resonance imaging. Radiology, 136: 707, 1980. 18. Schwartz, J. L. and Crooks, L. E.: NMR imaging produces no observable mutations or cytotoxicity in mammalian cells. Amer. J. Roentgen., 139: 583, 1982. 19. National Radiological Protection Board: Exposure to nuclear magnetic resonance clinical imaging. Radiography, 4 7: 258, 1981. 20. Davis, P. L., Crooks, L., Arakawa, M., McRee, R., Kaufman, L. and Margulis, A. R.: Potential hazards in NMR imaging: heating effects of changing magnetic fields and RF fields on small metallic implants. Amer. J. Roentgen., 137: 857, 1981. 21. Crooks, L., Arakawa, M., Hoenninger, J., Watts, J., McRee, R., Kaufman, L., Davis, P. L., Margulis, A. R. and DeGroot, J.: Nuclear magnetic resonance whole-body imager operating at 3.5 KGauss. Radiology, 143: 169, 1982. 22. Hricak, H., Crooks, L., Sheldon, P. and Kaufman, L.: Nuclear magnetic resonance imaging of the kidney. Radiology, 146: 425, 1983. 23. Hricak, H., Williams, R. D., Moon, K. L., Jr., Moss, A. A., Alpers, C., Crooks, L. E. and Kaufman, L.: Nuclear magnetic resonance imaging of the kidney: renal masses. Radiology, 147: 765, 1983. 24. Alfidi, R. J., Haaga, J. R., El-Yousef, S. J., Bryan, P. J., Fletcher, B. D., LiPuma, J. P., Morrison, S. C., Kaufman, B., Richey, J. B., Hinshaw, W. S., Kramer, D. M., Yeung, H. N., Cohen, A. M., Butler, H. E., Ament, A. E. and Lieberman, J. M.: Preliminary experimental results in humans and animals with a superconducting, whole-body, nuclear magnetic resonance scanner. Radiology, 143: 175, 1982.
25. Young, I. R., Bailes, D.R., Burl, M., Collins, A.G., Smith, D. T., McDonnell, M. J., Orr, J. S., Banks, L. M., Bydder, G. M., Greenspan, R. H. and Steiner, R. E.: Initial clinical evaluation of a whole body nuclear magnetic resonance (NMR) tomograph. J. Comput. Assist. Tomogr., 6: 1, 1982. 26. Hadley, M. D. M., Nichols, D. M. and Smith, F. W.: Nuclear magnetic resonance tomographic imaging in xanthogranulomatous pyelonephritis. J. Urol., 127: 301, 1982. 27. Pollet, J. E., Smith, F. W., Mallard, J. R., Ah-See, A. K. and Reid, A.: Whole-body nuclear magnetic resonance imaging: the first report of its use in surgical practice. Brit. J. Surg., 68: 493, 1981. 28. Newhouse, J. H.: Urinary tract imaging by nuclear magnetic resonance. Urol. Rad., 4: 171, 1982. 29. Smith, W. F., Hutchinson, J. M. S., Mallard, J. R., Johnson, G., Redpath, T. W. and Selbie, R. D.: Renal cyst or tumourdifferentiation by whole-body nuclear magnetic resonance imaging. Diagn. Imaging, 50: 61, 1981. 30. Smith, F. W., Reid, A., Hutchinson, J. M. S. and Mallard, J. R.: Nuclear magnetic resonance imaging of the pancreas. Radiology, 143: 677, 1982. 31. Hricak, H., Higgins, C. B. and Williams, R. D.: Nuclear magnetic resonance imaging in retroperitoneal fibrosis. Amer. J. Roentgen., 141: 35, 1983. 32. Hricak, H., Williams, R. D., Spring, D. B., Moon, K. L., Jr., Hedgcock, M. W., Watson, R. A. and Crooks, L. E.: Anatomy and pathology of the male pelvis by magnetic resonance imaging. Amer. J. Roentgen., 141: 1101, 1983. 33. Eggleston, J.C., Saryan, L.A. and Hollis, D. P.: Nuclear magnetic resonance investigations of human neoplastic and abnormal nonneoplastic tissues. Cancer Res., 35: 1326, 1975. 34. Davis, P. L., Crooks, L. E., Margulis, A. R. and Kaufman, L.: Nuclear magnetic resonance imaging: current capabilities. West. J. Med., 137: 290, 1982. 35. Brady, T. J., Goldman, M. R., Pykett, I. L., Buonanno, F. S., Kistler, J. P., Newhouse, J. H., Burt, C. T., Hinshaw, W. S. and Pohost, G. M.: Proton nuclear magnetic resonance imaging of regionally ischemic canine hearts: effect of paramagnetic proton signal enhancement. Radiology, 144: 343, 1982. 36. Young, I. R., Clarke, G. J., Bailes, D.R., Pennock, J.M., Doyle, F. H. and Bydder, G. M.: Enhancement of relaxation rate with paramagnetic contrast agents in NMR imaging. CT, 5: 543, 1981. 37. Brasch, R. C., London, D. A., Wesbey, G. E., Tozer, T. N., Nitecki, D. E., Williams, R. D., Doemeny, J., Tuck, L. D. and Lallemand, D. P.: Work in progress: nuclear magnetic resonance study of a paramagnetic nitroxide contrast agent for enhancement of renal structures in experimental animals. Radiology, 147: 773, 1983.
TBE JOURNAL 8F UROLOGY
Copyright© 1984 by The 'Williams & Wilkins Co.
Printed
October
U.S.A.
Review Article MAGNETIC RESONANCE IMAGING IN UROLOGY RICHARD D. WILLIAMS*
AND
HEDVIG HRICAK
From the Departments of Urology and Radiology, University of California School of Medicine and Veterans Administration Medical Center, San Francisco, California
Although still m its developmental stages, magnetic resonance imaging is expected to become one of the most significant advances in diagnostic body imaging. This optimistic view is based on the unique ability of magnetic resonance imaging not only to provide images of precise anatomical detail in virtually any planar projection desired but also to deliver simultaneously biochemical information concerning the tissues examined. Importantly, all information is obtained without imparting any known harmful effect on the subject, that is no ionizing radiation is used and iodinated contrast medium enhancement is unnecessary. The purpose of this review is to familiarize the urological surgeon with the underlying principles of magnetic resonance imaging and to present the current state of the art with respect to anatomical areas of interest to the urologist. MAGNETIC RESONANCE IMAGING-BASIC PHYSICAL PRINCIPLES
Even though the physical principles of magnetic resonance were described first in 19461 · 2 and the concept of magnetic resonance body imaging for medical use existed in the early 1970s3 - 8 success in adapting magnetic resonance for clinical use awaited development of the computers and software required for reconstruction of the i;;;l0 8 data bits necessary to produce x-ray computerized tomographic (CT) imaging. The fundamental concepts of magnetic resonance have been reported in detail recently and are based on the presence of atomic nuclei with an odd atomic number in living tissues. 8- 16 Examples of such atoms include 'hydrogen (1H), 13 carbon, 17 oxygen, 19fluorine (1 9F), 23 sodium (2 3 Na) and 31 phosphorus (31 P). Of these, hydrogen atoms (protons) are the most abundant based on the ubiquity of water (which contains 2 protons) and lipids within living tissues. Because these atoms have an electric charge and an intrinsic property of spin or angular momentum, each pro· duces a small magnetic field around it that can be described a vector (fig. 1, A). In undisturbed tissue the net magnetic vector of all protons present is zero (fig. 1, owing to the random orientation of the individual magnetic vectors. However, if a group of protons is placed in a strong external magnetic field, the magnetic vectors of the protons will tend to align in that field (fig. 1, C). Generally, they orient either parallel or anti-parallel to the external magnetic field. More protons will line up parallel (because this represents a lower energy state) and their net magnetic vector will be in the direction of the north pole of the external magnetic field. Magnetic field strengths are referred to in either Gauss or Tesla units (LO Tesla unit = 10,000 Gauss units). For example, the gravitational field of earth is approximately 0.6 Gauss units, whereas the field strengths used for magnetic resonance imaging are much stronger (0.1 to 1.5 Tesla units). Following alignment within the magnet each still-spinning proton begins to spin (precess) about the axis of the magnetic field, similar to a *Requests for reprints: Urology, U-518, University of California, San Francisco, California 94143.
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spinning top rotating around its own axis while rotating simultaneously around the vertical axis of the gravitational field of earth. The frequency of precession is characteristic for the external magnetic field strength. If energy in the radiofrequency range is applied at a right angle to the axis of the magnet at exactly the frequency of the protons' precession, their individual vectors merge because the added energy causes a rapid, repetitive unified flipping of the protons between the parallel and anti-parallel orientation (fig. 1, D). This change is termed resonance. The net magnetic vector of the protons will be displaced to a new position by the radiofrequency pulse. Because the protons absorb radiofrequency energy during this process, when the radiofrequency pulse ceases the absorbed radiofrequency energy is released as the protons' magnetic vectors return to their original state (fig. 1, E). The radiofrequency energy released will be emitted at exactly the same frequency as that of proton precession and can be detected by a radiofrequency wave monitor. It is this measurable release of energy that is called the magnetic resonance signal. The intensity of the magnetic resonance signal is proportional to the relative number of the atomic nuclei resonating. In living tissues only the atoms present in relative abundance can be used to provide a magnetic resonance image. 'H, 19F, 23 Na and 31 P always are present in these isotopic forms in living tissues, yet only 1 H is present in enough concentration and emits a strong enough signal for current use in clinical magnetic resonance imaging. Therefore, to date only proton magnetic resonance imaging has medical relevance. For medical magnetic resonance imaging we are more interested in the relative position or spatial relationship of 1 H atoms within the tissue being examined than merely their presence and relative numbers. To provide this information it is necessary to vary the strength of the external magnetic field, which is accomplished by 1m,rouu1.:1r1g another magnetic force within the aperture of the external. magnet so that a field gradient can be produced. Because the precessional frequency of each proton is related directly to its position in the graded external magnetic field the radiofrequency signal emitted by the protons following excitation contains spatial information. Therefore, by application of radiofrequency pulses of different duration and different frequency, the resultant magnetic resonance signals produced also will be slightly different and, when transformed and reconstructed by computer, can provide an image depicting the relative position, density and biochemical environment of the protons. Therefore, this information can reveal subtle differences within and between tissues. In fact, these differences can be measured more precisely than is possible in CT imaging and, therefore, can produce images with potentially greater anatomical discrimination. It is this variability in the field strength and, thus, excitation of selected protons at a given radiofrequency pulse within the tissue that allow image formation in almost any plane desired. Besides proton density images, at least 3 other magnetic resonance parameters can be used to impart different infor-
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A
B
-d3
C
N D
E
s
I
FIG. 1. A, hydrogen atom (proton) with atomic number of 1 results in angular momentum and magnetic vector. B, random distribution of protons in tissue with net magnetic moment of zero. C, proton magnetic vectors undergo alignment within strong external magnetic field. D, radiofrequency (RF) waves at right angles to magnetic field cause proton vector realignment (precession). E, discontinuation of radiofrequency waves permits proton vectors to return to original alignment in magnetic field, which releases radiofrequency energy (NMR, nuclear magnetic resonance signal), which then is processed by computer to produce image.
mation: Tl and T2 relaxation times, and bulk motion of hydrogen or flow. The first 2 parameters are based on measurements of proton precessional relaxation times. Tl refers to the time required for proton precession to return to thermal equilibrium as evidenced by the tendency for alignment (discussed previously) parallel to the external magnetic field after cessation of each radiofrequency pulse. This time can vary considerably: in solids at low temperatures, when random thermal motion is minimal, Tl may be long (hours), whereas in pure liquids Tl may be only several seconds. In liquids containing proteins (such as tissue fluids) Tl is even shorter. During image formation an interval (recovery time) must be introduced between successive radiofrequency pulses so that recovery of magnetization of protons after the radiofrequency pulse ceases can occur and be measured. Increasing recovery time can increase the intensity of the magnetic resonance signal but also can increase the scanning time. Therefore, recovery time must be selected carefully for medical imaging. T2 relaxation time refers to the signal decay time induced by the magnetic forces exerted on the protons in their microenvironment (affected by adjacent protons), which dampen coherent resonance until the signal ceases. T2 is measured in milliseconds and Tl in seconds. The interval between an applied radiofrequency pulse and signal reception is a measure of the role of signal decay and is designated by the instrument parameter (echo delay time). In solids T2 is in the microsecond range, while in liquids T2 is longer. Thus, variable instrument setting and the salient tissue characteristics affect the magnetic resonance signal intensity and images can be reconstructed, which relate to proton density, and Tl and T2 relaxation times. Tl and T2 images are grayscale maps synthesized from the different combinations of recovery time and echo delay time used to sample intensity of
the magnetic resonance signal at different points in time. Since approximately 50 msec. are needed for the proton signal to register, if the protons move faster through the plane of imaging, their signal will not be detected. This permits assessment of blood flow and represents the third magnetic resonance parameter, proton bulk motion. Currently, there are 4 major techniques used to obtain magnetic resonance intensity images: 1) partial saturation recovery, 2) free inductive decay, 3) spin echo and 4) inversion recovery. At our university spin echo and inversion recovery data acquisition techniques are used routinely. In the spin echo technique the image is dependent on the Tl and T2 relaxation parameters, hydrogen density (spin density) and blood flow. The instrument parameters that can be varied to produce tissue contrast with the spin echo technique are recovery time and echo delay time. Recovery time represents the interval between the repetitive radiofrequency pulses applied to the image volume and it can be varied from 0.5 to 2 seconds. Echo delay time indicates the interval from excitation of the nuclei by radiofrequency pulse and the receipt of the signal from the nuclei. Images are obtained at echo delays of 28 and 56 msec. in each section plane. In the inversion recovery technique the image is dependent on hydrogen density and Tl relaxation values. The instrument parameters that can be varied to produce tissue contrast are recovery time, inversion time and, at our institution, echo delay time as well. Recovery time is referred to as the interval between repetitive 90-degree radiofrequency pulses applied to image volume and it can be 1 or 1.8 seconds. Inversion time is the interval between initiation of 180 and 90-degree pulses, and it can be 420 or 277 msec. Echo delay time represents the interval between excitation of the nuclei by radiofrequency pulse and receipt of the signal from the nuclei. On inversion recovery, similar to spin echo, images are obtained at echo delay times of 28 and 56 msec. The necessary components of a magnetic resonance imager are a magnet capable of imposing a strong uniform magnetic field, gradient coils that can alter the magnetic field internally depending on the technique used, a transmitter to deliver radiofrequency waves to the subject, a receiver to gather magnetic resonance signals from the subject and deliver them to appropriate digitizing circuitry and, finally, a computer capable of processing the information received into a spatial display (fig. 2). Three different types of magnets are available: 1) permanent, 2) resistive and 3) superconducting. Currently, 2 major approaches provide magnets of sufficient strength for medical magnetic resonance imaging. The resistive coil electromagnet can produce useful images at a reasonable cost (approximately $800,000). However, the homogeneity of the field produced in this device (an important factor for image accuracy) cannot be greater than a few parts in 105 and, because the power requirements and thermal dissipation needs for this type of magnet are great, the maximal field strengths reachable range from 1 to 2 KGauss units (0.2 Tesla units). The superconductive magnet has the advantages of greater obtainable field strength (theoretically up to 14 Tesla units), superior homogeneity (1 part in 108 ) and increased stability. The basic magnet consists of wound filaments of a superconductor that are supercooled by liquid nitrogen and helium. Electricity is required to initiate the magnet but in the supercooled environment the superconductor has almost no resistance and, therefore, can conduct and maintain the current necessary to produce the magnetic field indefinitely. Therefore, the power requirements of the magnet are minimal. However, the initial cost of the superconductive magnet is approximately double that of the resistive device. Because superconductive magnets produce a field stronger and more stable than that of resistive magnets, and because possible imaging of nuclei other than protons will require a
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MAGNETIC RESONANCE IMAGING IN UROLOGY
A
B
~,,o
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Computer
0
Image Display System
D
FIG. 2. Typical superconductive magnetic resonance imager. A, external view. B, longitudinal schematic view of components necessary for magnetic resonance image construction.
strong magnet in the future, the superconductive magnet at present has the preferential advantage. ADVANTAGES OF MAGNETIC RESONANCE IMAGING
The enthusiasm attending the emergence of magnetic resonance imaging for medical use is predicated upon certain advantages over current conventional radiological imaging techniques that already have been established: 1) magnetic resonance imaging avoids the use of ionizing radiation, 2) magnetization in the power range used is without harmful biological effects, 3) spatial resolution comparable to the latest generation x-ray CT imaging already is demonstrable, 4) contrast resolution is vastly superior, 5) iodinated contrast medium is not required, 6) no beam hardening or bone artifact (particularly important in imaging the pelvis) is seen, 7) blood flow measurement is possible, 8) imaging of 31 P, 19F or 23 Na may permit determination of cellular or tissue biochemical abnormalities and 9) image distortion owing to metal objects (such as surgical clips) is absent. The development of any new diagnostic modality also must be scrutinized with respect to possible risks, toxicity or disadvantages. The initial cost of the magnet, slow scanning time (compared to CT), and lack of detection of cortical bone and tissue calcium appear to be the only deterrents of magnetic resonance imaging. Similar to the early problems with CT, claustrophobia has been a problem in approximately 5 per cent of the patients examined. In vitro cellular toxicity studies have revealed no evidence of inhibition of deoxyribonucleic acid (DNA) metabolism and DNA function remains unaffected. 17• 18 Human studies have shown no short-term or medium-term effects despite multiple exposures of subjects and operator personnel. 19 The experience in our laboratory has shown no observed toxicity in > 1,000 patients with a wide variety of clinical conditions, thereby eliminating many of the limitations reported previously. 20 • 21 CLINICAL MAGNETIC RESONANCE IMAGING
The images presented were obtained on our magnetic resonance imager, a superconducting electromagnet functioning at
a field strength of 0.35 Tesla units (15 MHz.). 21 Images were produced via a selective irradiation technique with 2-dimensional Fourier transformation. The imaging sequences chosen in our laboratory were spin echo and inversion recovery. Spin echo is accomplished using radiofrequency repetition intervals of 500 to 2,000 msec., and echo delay times of 28 and 56 msec. These parameters allow as many as 20 anatomic sections to be obtained simultaneously in approximately 18 minutes. Inversion recovery is accomplished with repetition intervals of 1,000 and 1,800 msec., with an inversion to recovery of 420 and 280 msec., respectively, allowing up to 9 sections in 16 minutes. The choice of imaging sequences is important, since time limitations dictate that only 2 to 3 sequences can be accomplished at any one sitting to keep scanning time between 30 and 60 minutes. Because the art is evolving, the precise combinations of the variables required to produce discriminating images in each clinical situation have yet to be determined. MAGNETIC RESONANCE IMAGING GRAY SCALE
Magnetic resonance imaging is a function of mobile protons (spin density), Tl and T2 relaxation parameters, and bulk motion of protons (flow). Contrast on the magnetic resonance image is a relative phenomenon and is the result of inherent tissue characteristics, the imaging technique used (spin echo, free inductive decay, partial saturation recovery and inversion recovery) and the magnetic field strength used. Within each imaging technique tissue contrast is influenced by instrument parameters (recovery time, echo delay time and inversion recovery time). The gray scale in the spin echo magnetic resonance image shows fat as the brightest tissue followed by marrow and cancellous bone, brain and spinal cord, viscera, striated muscle, ligaments and tendons, rapidly flowing blood, compact bone and air in decreasing order of intensity. With the spin echo technique the signal acquisition, variations of echo delay time and recovery time will discriminate between tissue as a function of their relative length of relaxation times. The tissue contrast seen with the inversion recovery technique varies with the inversion recovery interval and recovery time. This technique
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FIG. 3. Normal kidneys. A, spin echo recovery time is 1 second and echo delay time is 28 msec. There is clear differentiation between higher intensity cortex ( C) and lower intensity medulla (M). Gerota's fascia (large arrow) is seen as line dividing perirenal from pararenal spaces. Small arrow denotes left renal vein. A, aorta. I, inferior vena cava. L, liver. B, inversion recovery. Differentiation between lower intensity medulla (M) and higher intensity cortex ( C) is more pronounced than on spin echo image. L, liver.
emphasizes variations in Tl relaxation time, while the spin echo technique is dependent on Tl and T2 relaxation times. NORMAL RENAL ANATOMY
The anatomic features of a normal kidney are defined clearly by magnetic resonance imaging (fig. 3). On spin echo images the renal hilus is imaged as a high intensity area owing to the presence of hilar adipose tissue. The renal veins and arteries are imaged as low intensity tubular structures traversing the hilar region. The low intensity of the vessel lumen is caused by rapid blood flow. The pelviocaliceal system and the ureters are imaged with low intensity owing to the long Tl and T2 relaxation values of urine. On images obtained during diuresis a mild dilatation of the pelviocaliceal structures can be appreciated. Within the renal parenchyma the renal cortex is distinguished easily from the medulla, since the cortex is of a higher intensity than the medulla yet lower than the perirenal fat (fig. 3, A). The medulla is a triangular low intensity area central to the cortex that has an intensity and prominence related to the state of hydration. 22 This fact suggests that imaging will be better with the patient in a hydrated rather than dehydrated state. The cortical medullary differentiation is enhanced on short recovery time pulse sequences and virtually disappears on images obtained with a long recovery time (2 seconds). Perirenal fat is seen as a high intensity area owing to the short Tl and long T2 relaxation times of adipose tissue. Gerota's fascia is visible readily as a low intensity line surrounding the perirenal space. On inversion recovery images the renal cortex is significantly brighter than the medulla and differentiation between the two is more evident than on spin echo images (fig. 3, B). Perirenal and hilar fat remains bright owing to a short Tl relaxation time. Urine remains dark because of the long Tl relaxation time (fig. 3, B). RENAL PATHOLOGICAL FINDINGS
FIG. 4. A, left ureteropelvic junction obstruction. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Right kidney is normal with excellent cortical medullary differentiation. Arrow shows normal right ureter. Calices ( C) of left kidney are dilated and there also is marked dilatation of left extrarenal pelvis (P). Cortical medullary differentiation of left kidney is not present. Note thickened Gerota's fascia (curved arrow). B, chronic glomerulonephritis. Both kidneys are small with marked thickening of renal parenchyma ( C). Simple renal cyst (arrow) arises from posterolateral aspect of right kidney. Note abundant hilar adipose tissue (asterisk).
Hydronephrosis is detectable easily by magnetic resonance imaging, as is a hydroureter. By varying spin echo instrument parameters the distended ureter can be differentiated from adjacent bowel, lymph nodes or blood vessels. On CT images contrast enhancement is required to make this differentiation. An example of hydronephrosis on magnetic resonance imaging is shown in figure 4, A. Because fat and fluid have opposite magnetic resonance imaging intensity signals, peripelvic cysts and sinus lipomatosis are distinguished readily by magnetic resonance imaging, whereas partial averaging on CT can cause ambiguity. In the diagnosis of renal parenchymal diseases magnetic resonance imaging promises to make a significant contribution (fig. 4, B). The increased sensitivity of magnetic resonance imaging to subtle changes in soft tissues suggests that monitoring the progression of medical renal disorders may be possible,
perhaps limiting the need for renal biopsy. In addition, because acquired renal cystic disease and renal neoplasms are increased in patients with chronic renal failure, magnetic resonance imaging, obviating the use of iodinated contrast agents, may prove efficacious. The ability of magnetic resonance imaging to distinguish solid from cystic renal lesions is accurate and comparable to that of contrast-enhanced CT. Cysts are of homogeneous low intensity and without internal structures (fig. 5). Tl values of cysts usually are prolonged and T2 values are >200 msec. A smooth interface with renal parenchyma usually is demonstrated. Cyst walls are not demonstrated. Hemorrhagic and infected cysts have been shown to have an increased spin echo intensity compared to simple cysts, suggesting that magnetic resonance imaging has a diagnostic advantage over other modalities in
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the evaluation of patients with renal trauma, or infected or hemorrhagic cysts (fig. Evaluation of renal tumors has shown a wide spectrum of magnetic resonance imaging findings (fig. 7, A and B). Spin echo images show a range from low to very high signal intensity lesions. 23 · 24 Prolongation of the Tl relaxation time of solid tumors relative to adjacent normal renal parenchyma is a common feature, whereas the T2 values are similar or increased relative to surrounding parenchyma. Staging of renal neoplasms by magnetic resonance imaging appears particularly promising in the evaluation of neoplastic involvement of the renal vein or inferior vena cava (fig. 7, C). A potential disadvantage of magnetic resonance imaging in the evaluation of renal masses is the lack of definition of calcification despite a variety of magnetic resonance imaging techniques. Calcifications have low intensity and frequently are not visualized. When seen, calcifications are not differentiated readily from dense fibrous tissue owing to a similar low intensity signal. Whether this feature of magnetic resonance imaging will prove to be a significant weakness requires further study. Although limited information is available to date it is expected that further experience will reveal magnetic resonance imaging to be of substantial value in defining renal parenchymal disease, detecting and categorizing renal masses, and staging renal neoplasms.
FIG. 5. Simple renal cyst. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Homogeneous low intensity cyst (C) is in right kidney. Cyst wall is not demonstrated but there is sharp demarcation at border with normal renal parenchyma (arrow). A, aorta. I, inferior vena cava. SMA, superior mesenteric artery. SMV, superior mesenteric vein. L, liver. S, spleen.
FIG. 7. A, renal cell carcinoma (T) in upper pole of right kidney is imaged with inhomogeneous high intensity. Low intensity line (arrow) interposed between tumor and renal parenchyma represents tumor pseudocapsule. L, liver. S, spleen. B, renal cell carcinoma is imaged as inhomogeneous predominantly low intensity lesion ( T) arising from medial aspect of right kidney. Arrow shows left renal vein. C, tumor thrombus (T) within inferior vena cava is imaged with medium intensity signal and is seen occupying entire lumen of inferior vena cava. Arrow shows left renal vein. Curved arrows show right and left renal arteries. There is hydronephrosis of right kidney. NORMAL ADRENAL ANATOMY
The adrenal gland is homogeneous and of low intensity on magnetic resonance imaging, and is surrounded and outlined clearly by the higher intensity periadrenal fat (fig. 8). The left adrenal is visualized routinely but the right adrenal is seen slightly less often. The left adrenal is an inverted V or a triangle, whereas the right gland appears superiorly as a thin line extending from the inferior vena cava juxtaposed between the ems of the diaphragm and the medial aspect of the right lobe of the liver, and inferiorly as a horizontal band posterior to the inferior vena cava. Differentiation of the adrenal cortex and medulla is possible, with the medulla being of lower intensity than the cortex. Subjectively, the adrenal medulla has longer Tl and slightly shorter T2 relaxation times than the outer rim. ADRENAL PATHOLOGICAL FINDINGS
FIG. 6. Polycystic kidney disease. Spin echo recovery time is 2 seconds and echo delay time is 28 msec. Both kidneys are enlarged markedly. Renal parenchyma is replaced by renal cysts of various size and intensity. Acutely hemorrhagic cyst is seen with high intensity (arrow).
Primary as well as metastatic adrenal lesions are detected easily by magnetic resonance imaging (fig. 9, A). Currently, differentiation between malignant and benign lesions is not possible. Since the size and volume of the gland can be dis-
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cerned adrenal hyperplasia is displayed easily by magnetic resonance imaging (fig. 9, B). These early results suggest that magnetic resonance imaging can match the ability of CT to demonstrate normal and pathological adrenal anatomy. Because of the ability of magnetic resonance imaging to differentiate adrenal cortex and medulla, and easy anatomical display of the adrenals (surrounding fat, and the adjacent liver and large blood vessels), magnetic resonance imaging eventually may surpass CT in the differential diagnostic evaluation of the adrenal gland.
nance imaging showed that the tumor was separate from the kidney and that the renal vein was patent. The leiomyosarcoma was inhomogeneous and contained a large high intensity central area documented at operation to be a liquefied hematoma. The proximal extent of the vena caval thrombus was capped by a high intensity lesion that proved to be a clot on the leading edge of the tumor. Collateral venous blood vessels coursing around the thrombosed vena cava were seen easily (fig. 11). These features were not appreciated even in retrospect by the CT studies.
RETRO PERITONEUM
Little experience has been gained to date in magnetic resonance imaging of the retroperitoneum other than that described previously in relation to renal and adrenal anatomy. 25 - 30 Retroperitoneal fibrosis is imaged as an irregularly shaped medium to low intensity mass. 31 The signal intensity is higher in the plaques than in adjacent muscle but less than that of adjacent fat. Tl values varied considerably but T2 values were low consistently. The extent of involvement is determined easily by magnetic resonance imaging and is striking on sagittal scans. Because of the lack of signal from vessels with flowing blood, the perivascular component and consequent vessel narrowing are demonstrated easily. Evaluation of the vascular involvement by retroperitoneal fibrosis is superior by magnetic resonance imaging than contrast-enhanced CT (fig. 10). We also have examined 1 case of a retroperitoneal leiomyosarcoma occupying the entire right retroperitoneal space with an occlusive tumor thrombus in the inferior vena cava. Magnetic reso-
FIG. 8. Normal adrenal gland. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Lateral and medial limbs of left adrenal gland (arrow) are imaged clearly. Within kidneys there is superb differentiation between lower intensity medulla (M) and higher intensity cortex (C). St, food-filled stomach. L, liver.
FIG. 9. A, adrenal metastases from carcinoid tumor (inversion recovery image) are imaged as low intensity lesions (arrow). Large metastatic focus (small arrows) is within liver (L). St, stomach. S, spleen. B, adrenal hyperplasia. Spin echo image recovery time is 1 second and echo delay time is 28 msec. Both adrenal glands are enlarged (arrows).
FIG. 10. Retroperitoneal fibrosis. A, magnetic resonance imaging spin echo image recovery time is 1 second and echo delay time is 28 msec. Retroperitoneal plaque (P) encircles aorta. Normal flow through aorta (arrow) is demonstrated. Curved arrow shows inferior vena cava. Arrowhead shows stent within left ureter. B, CT scan shows retroperitoneal plaque encircling great vessels (P). Calcifications are within aortic wall. No information regarding flow is obtained. Inferior vena cava cannot be separated from plaque. Arrowhead shows stent within ureter.
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and urine is determined easily, and wall thickness and irregularity are appreciated readily. 33 Pelvic bones are distinguished easily by magnetic resonance imaging. Bone marrow is visualized with a high intensity signal. The surrounding cortical bone emits no signal owing to lack of mobile protons. If the bladder is not distended bowel loops can present confusing artifacts at the dome of the bladder. However, analysis of several anatomic sections with a variety of pulse sequence intervals will permit the proper distinction to be made. BLADDER PATHOLOGICAL FINDINGS
FIG. IL Leiomyosarcoma of inferior vena cava. Large right retroperitoneal mass has inhomogeneous intensity. High intensity area (H) represents liquefied hematoma. Tumor (T) invades entire vena cava.
Hypertrophied bladder wall has a medium intensity signal (fig. 13). Bladder tumors have a signal intensity higher than bladder muscle and, thus, can be distinguished easily from the normal bladder wall (fig. 14, A and B). Depth of penetration of bladder neoplasms is appreciated best by combining axial and sagittal images. Magnetic resonance imaging appears to show particular promise with respect to local staging. Also, magnetic resonance imaging is superior to CT in the evaluation of tumors located at the bladder base. NORMAL PROSTATE ANATOMY
The normal prostate reveals a homogeneous medium intensity signal on magnetic resonance imaging (fig. 14, C). With multiplanar images, contiguous structures, such as bladder base, rectum and periprostatic fat, readily are apparent. The seminal vesicles have a medium intensity signal and are clearly distinguished embedded within the pelvic fat (fig. 15). PROSTATE PATHOLOGICAL FINDINGS
Benign prostatic hyperplasia is imaged with a homogeneous medium intensity signal (fig. 16, A). Prostatic carcinoma appears to have a higher intensity signal than normal or adenomatous tissue (fig. 16, B). Involvement of the seminal vesicles (stage C), confirmed by palpation and pathological examination, is evident on magnetic resonance imaging. Distortion of the prostatic contour and disruption of the lower intensity surgical capsule are seen in patients with stage C disease. If these findings are confirmed when larger numbers of patients are studied magnetic resonance imaging will emerge as the modality of choice in the diagnosis and local staging ofprostatic neoplasms. In addition to the anatomic areas described, there is a superb display of penile and inguinal canal anatomy, with differential
FIG. 12. Sagittal scan of male pelvis shows Denonvilliers' fascia (arrow) separating prostate gland (P) from rectum (R). B, bladder. SV, seminal vesicles. CS, corpora spongiosa. CC, corpora cavernosa. SP, symphysis pubis.
PELVIS
Lack of respiratory motion, no bone or beam hardening artifact and superior contrast resolution make magnetic resonance imaging the superior modality in the evaluation of the pelvis. 32 NORMAL VESICAL ANATOMY
The bladder is distended routinely with urine during scanning and, therefore, is demonstrated easily (fig. 12). The signal intensity of urine is low. 32 Differentiation between vesical wall
FIG. 13. Bladder wall hypertrophy. Bladder wall (arrows) is thickened markedly and of homogeneous medium intensity signal that increases with longer recovery times. A, spin echo recovery time is 1 second and echo delay time is 28 msec. B, recovery time is 2 seconds and echo delay time is 28 msec. R, rectum.
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FIG. 15. Seminal vesicles (SV) are imaged with medium intensity signal posterior to bladder (B). R, rectum.
FIG. 14. A, transitional cell carcinoma ( T) is seen along right lateral and posterior walls of bladder (small arrows). Bulk of tumor protrudes within bladder lumen. Tumor also involves dilated right ureter (curved arrow). R, rectum. B, adenocarcinoma of urachal cyst (T) is imaged with medium intensity at dome of bladder. High intensity lesion outlining bladder lumen represents transitional cell carcinoma (asterisks). Foley catheter (arrow) is seen within urethra. P, prostate gland. SV, seminal vesicles. R, rectum. C, axial scan shows normal prostate (P) imaged with medium intensity signal. Small arrows show levator ani muscle. R, rectum.
demonstration of the cord structures, and excellent images of the testes. Pathological conditions of these areas have yet to be imaged. Magnetic resonance imaging of the pelvis appears to have contrast resolution capabilities exceeding that of CT, even with the use of contrast enhancement. Staging of pelvic neoplasms appears to be improved by the ability to examine the pelvis in multiple planes. Currently, the substantial obstacle appears to be the inability to depict pelvic lymph nodes clearly. However, this problem is expected to be resolved with experience and improvement in scanning techniques. FUTURE PERSPECTIVES
Our view of the potential of magnetic resonance imaging is optimistic. To date, the science is in its infancy. Nevertheless, for many organ sites the images produced by early generation magnetic resonance imagers already are equal in anatomical
resolution to those produced by current generation CT scanners. Furthermore, magnetic resonance images are superior to enhanced CT in other organs, most notably the brain, retroperitoneum and pelvis. With the advent of improved pulse sequencing techniques, perhaps specific for organs or disease states, the art should advance rapidly. In the near future it should be possible to obtain blood and urine flow measurements quantitatively, determine presence and volume of muscle ischemia or infarction and extend magnetic resonance imaging to other atoms, such as 31 P or 19F, which may allow specific physiological information to be gained. Perhaps a more exciting recent development in magnetic resonance imaging is the use of contrast agents to enhance the anatomic information now available. These agents are paramagnetic and, thus, shorten the Tl and T2 relaxation times. 34- 36 Manganese chloride already has been used to reduce magnetic resonance signal intensity from zones of myocardial infarction in isolated perfused canine hearts, 35 oral ferric chloride has been used successfully as an intraluminal bowel contrast agent, and inhaled 100 per cent oxygen has been used as an intracardiac blood magnetic resonance signal enhancer in human volunteers. Recently, nitroxide stable free radicals have been used effectively as an intravenous urographic contrast agent in animals. 37 If these developments are applicable routinely in humans, magnetic resonance imaging may well surpass currently available diagnostic modalities for selected organs and disease states. The ultimate role of magnetic resonance imaging in diagnostic medicine will require considerable research for definite conclusions to be drawn. However, in urology magnetic resonance imaging already has begun to establish its superiority in distinct anatomical sites and disease states. REFERENCES 1. Bloch, F.: Nuclear induction. Phys. Rev., 70: 460, 1946. 2. Purcell, E. M., Torrey, H. C. and Pound, R. V.: Resonance absorption by nuclear magnetic moments in a solid. Phys. Rev., 69: 37, 1946. 3. Jackson, J. A. and Langham, W. H.: Whole-body NMR spectrometer. Rev. Sci. lnstrum., 39: 510, 1968. 4. Damadian, R.: Tumor detection by nuclear magnetic resonance. Science, 171: 1151, 1971. 5. Lauterbur, P. C.: Image formation by induced local interactions: examples employing nuclear magnetic resonance. Nature, 242: 190, 1973. 6. Lauterbur, P. C.: Magnetic resonance zeugmatography. Pure Appl. Chem., 40: 149, 1974. 7. Hinshaw, W. S., Bottomley, P.A. and Holland, G. N.: Radiographic thin section image of the human wrist by nuclear magnetic resonance. Nature, 270: 722, 1977. 8. Pykett, I. L., Newhouse, J. H., Buonanno, F. S., Brady, T. J., Goldman, M. R., Kistler, J.P. and Pohost, G. M.: Principles of nuclear magnetic resonance imaging. Radiology, 143: 157, 1982. 9. Pykett, I. L.: NMR imaging in medicine. Sci. Amer., 246: 78, 1982. 10. Oldendorf, W. H.: NMR imaging: its potential clinical impact.
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