Surgery – Diagnostic testing – Detecting nuclear – electromagnetic – or ultrasonic radiation
Reexamination Certificate
2002-03-28
2003-11-18
Smith, Ruth S. (Department: 3737)
Surgery
Diagnostic testing
Detecting nuclear, electromagnetic, or ultrasonic radiation
C324S318000
Reexamination Certificate
active
06650926
ABSTRACT:
BACKGROUND OF THE INVENTION
The present invention relates to magnetic resonance imaging (MRI) systems and, in particular, to the radio frequency (RF) coils used in such systems.
Magnetic resonance imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in a volume of interest (for example, in the human body), which are polarized by a strong, uniform, static magnetic field of the magnet (named B
0
—the main magnetic field in MRI physics). The magnetically polarized nuclear spins generate magnetic moments in the volume of interest. The magnetic moments point in the direction of the main magnetic field in a steady state, and produce no useful information if they are not disturbed by any excitation.
The generation of nuclear magnetic resonance (NMR) signals for MRI data acquisition is accomplished by exciting the magnetic moments with a uniform RF magnetic field (B
1
field or the excitation field). The B
1
field is produced in the imaging region of interest by an RF transmit coil which is driven by a computer-controlled RF transmitter with a power amplifier. During excitation, the nuclear spin system absorbs magnetic energy, and its magnetic moments precess around the direction of the main magnetic field (B
0
). After excitation, the precessing magnetic moments will go through a process of free induction decay (FID), releasing their absorbed energy and returning to the steady state. During FID, NMR signals are detected by the use of a receive RF coil, which is placed in the vicinity of the excited volume of interest. The NMR signal is the secondary electrical voltage (or current) in the receive RF coil that has been induced by the precessing magnetic moments of the human tissue or other volume of interest. The receive RF coil can be either the transmit coil itself, or an independent receive-only RF coil. The NMR signal is used for producing magnetic resonance images by using additional pulsed magnetic gradient fields, which are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of interest. There are usually three sets of gradient coils in a standard MRI system, which generate magnetic fields in the same direction of the main magnetic field, varying linearly in the imaging volume.
In MRI, it is desirable for the excitation and reception to spatially uniform in the imaging volume for better image uniformity. In a standard MRI system, the best excitation field homogeneity is usually obtained by using a “whole-body” volume RF coil for transmission. The “whole-body” transmit coil is the largest RF coil in the system. A large coil, however, produces lower signal-to-noise ratio (SNR or S/N) if it is also used for reception, mainly because of its greater distance from the signal-generating tissues being imaged. Since a high signal-to-noise ratio is desirable in MRI, special-purpose coils are used for reception to enhance the S/N ratio from the volume of interest.
In practice, a well-designed specialty RF coil is has the following functional properties: high S/N ratio, good uniformity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the coil device should be mechanically designed to facilitate patient handling and comfort, and to provide a protective barrier between the patient and the RF electronics. Another way to increase the SNR is by quadrature reception. In this method, NMR signals are detected in two orthogonal directions, which are in the transverse plane or perpendicular to the main magnetic field. The two signals are detected by two independent individual coils which cover the same volume of interest. With quadrature reception, the SNR can be increased by a factor of up to the square root of 2 over that of the individual linear coils.
In MRI, a torso-pelvis RF coil is used to image the human torso region from the top of the liver to the iliac crest and the pelvic region from the iliac crest to the pubic symphysis. Abdominal and pelvic imaging a torso-pelvis coil to be able to provide good image uniformity in the axial direction (i.e., the transverse direction) as well as a good SNR. Non-uniform images caused by a inhomogeneous signal sensitivity profile of a RF coil can lead to misdiagnosis of patients. For example, a high signal in the anterior abdomen region can be mistaken for a peritoneal tumor. The coverage of a torso-pelvis coil needs to be about 30 cm for the torso imaging and about 28 cm for the pelvic imaging, respectively, for most of the patient population. A torso-pelvis coil is also required to be able to image the entire torso and pelvic regions, i.e., to cover the entire 58 cm field-of-view (FOV), without repositioning a patient. So far, the usable imaging volume, which is defined as a sphere, of a commercially available MRI apparatus is about 45 to 50 cm in diameter. Therefore, two consecutive scans are needed to cover the 58 cm FOV.
The built-in birdcage (Hayes, U.S. Pat. No. 4,692,705) transmit and receive “whole-body” coil of most MRI scanners can be used to image a patient's abdomen and pelvis with good image homogeneity and without repositioning the patient. However, the major drawback of using a “whole-body” coil as a receive coil is that the SNR is too low. The low SNR of a “whole-body” coil is caused by the low filling factor and also by the noise/unwanted signals from the tissue outside the region-of-interest (ROI). The filling factor of a RF coil is determined by the ratio of the volume of the sample (e.g., a human patient's body) being imaged to the effective imaging volume of the coil. The closer the filling factor of a RF coil to unity, the better SNR of the coil. Usually, a “whole-body” coil has an effective imaging volume much bigger than the volume of the body portion of a patient being imaged. A “whole-body” coil most the time covers a much bigger FOV (about 48 cm) than the body portion of interest in imaging (e.g., 30 cm for the torso imaging). This causes the “whole-body” coil to couple to more noise and unwanted signals from the tissue outside the ROI or volume of interest and thus to have a lower SNR.
To increase the filling factor, an elliptical birdcage coil was developed (Leifer, U.S. Pat. No. 5,986,454). In this design, the circular cross-section of a conventional birdcage coil was modified into an elliptical one so that it provides a higher filling factor for imaging the human body portion, for example: the torso, whose cross-section is more elliptical than circular. The current distribution on each of the legs of the elliptical birdcage coil was also optimized in order to obtain a homogeneous image. This elliptical birdcage coil was operated in quadrature mode.
The development of array coil technology (Roemer, et al., U.S. Pat. No. 4,825,162) allows one to image a large field-of-view while maintaining the SNR characteristic of a small and conformal coil. Using this concept, a four element “C-shaped” adjustable volume array coil was built (Jones, U.S. Pat. No. 5,477,146) to improve the SNR for volume imaging. The mechanical housing of the “C-shaped” volume array coil is divided into two parts: anterior and posterior. Electrically, the “C-shaped” volume array coil consists of four loop coils: three loop coils in the anterior housing and one loop coil in the posterior housing. Each loop coil is critically coupled to its adjacent coil or coils to minimize the inductive coupling between the two adjacent coils and hence to reduce the noise correlation caused by the cross-talk between them. The anterior housing is pivotally connected to the posterior housing by a hinge at the location close to the posterior base. This allows the anterior housing to be adjusted for positioning a patient and for fitting different size patient into the coil as well.
It has been realized that the direction of the magnetic field generated by a butterfly coil (or saddle coil) can be perpendicular to that generated by a loop coil. Thus, by using a
Chan Pei Hsuon
Zou Mark Xueming
Smith Ruth S.
USA Instruments, Inc.
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