Abstract: A magnetic resonance imaging machine includes a toroidal vacuum dewer (24) which contains a superconducting magnet (10). A radio frequency coil (32) is mounted within a cylindrical bore (26) of the vacuum dewer. A cylindrical, dielectric former (46) supports an RF shield (34), a z-gradient coil (50), an x-gradient coil (52), and a y-gradient coil (54). The x and y-gradient coils are each composed of four like spiral coil constructions. A metallic layer is cut with cut lines (64) to define a generally spiral coil winding pattern. In a high current density region (68) in which the coil windings are narrower than a preselected width, the cut lines (76) are thinner. In lower current density regions (70), the cut lines (78) are thicker. In lower current density regions, two cut lines are defined between adjacent coil windings such that the coil windings are limited to a maximum width.

Abstract: The magnetic field assembly of a magnetic resonance imaging device includes an annular superconducting magnet (10) which is mounted within a toroidal vacuum vessel (24). A cylindrical member (26) defines a central bore through which the superconducting magnets generate a temporally constant primary magnetic field. A cylindrical, dielectric former (46) is mounted in the bore displaced a small distance from the cylindrical member. A radio frequency coil (32) is mounted within the cylindrical member defining a patient receiving examination region. An RF shield (34) is mounted around the exterior peripheral surface of the former. Primary gradient coils (40) are mounted around and potted to the exterior of the dielectric former around the RF shield. Gradient shield or secondary coils (44) are potted around an exterior of the cylindrical member within the vacuum chamber. As illustrated in FIG.

Abstract: The magnetic field assembly of a magnetic resonance imaging device includes an annular superconducting magnet (10) which is mounted within a toroidal vacuum vessel (24). A cylindrical member (26) defines a central bore (12) through which the superconducting magnets generate a uniform, static magnetic field. A cylindrical, dielectric former (46) is mounted in the bore displaced by an annular gap (58) from the cylindrical member. A shimset (60) for shimming the uniformity of the magnetic field is mounted in the gap (58). A radio frequency coil (32) is mounted within the cylindrical member defining a patient receiving examination region. An RF shield (34) is mounted around the exterior peripheral surface of the former. Primary gradient coils (50, 52, 54) are mounted around and potted to the exterior of the dielectric former around the RF shield. Gradient shield or secondary coils (74, 76, 78) are potted around an exterior of the cylindrical member within the vacuum chamber.

Abstract: A superconducting magnetic imaging apparatus includes a vacuum vessel (40) having a central helium reservoir (48) in which superconducting magnetic coil windings (44) are maintained at a superconducting temperature. The vacuum vessel defines an internal bore (42) within which a self-shielded gradient coil assembly (14) and an RF coil (22) are received. The self-shielded coil assembly includes an inner former (60) which defines an imaging region (12) within which an imaged portion of the patient are received. X and y-gradient coils having winding patterns (62) are bonded to the former (60) forming an integral structure. A z-gradient coil (70) is mounted to mechanical reinforcement structure (72) to be held in a spaced relationship from the x and y-gradient coils with an air gap (74) in between. This facilitates the dissipation of heat generated by the large current pulses applied to the x and y-gradient coils.

Abstract: An examination region (12) is defined within the bore of a superconducting magnet assembly (10). An RF coil (22) and gradient magnetic field coils (14) are disposed within the bore of the superconducting magnet assembly around the examination region. The superconducting magnet includes a hollow, cylindrical vacuum vessel (40). An annular, liquid helium holding low temperature reservoir (60) extends centrally through the vacuum vessel, but is sealed therefrom such that liquid helium is not drawn into the vacuum. A plurality of annular superconducting magnets (56) are received in the low temperature reservoir immersed in the liquid helium. A first cold shield (44) and a second cold shield (50) are mounted in the vacuum vessel surrounding the low temperature reservoir. A main magnetic field shield coil (66) is disposed in the low temperature reservoir outside of the annular superconducting magnets for canceling the magnetic field generated by the annular magnets surrounding the magnet.

Abstract: An examination region (12) is defined within the bore of a superconducting magnet assembly (10). An RF coil (22) and gradient magnetic field coils (14) are disposed within the bore of the superconducting magnetic assembly around the examination region. The superconducting magnet includes a hollow, tubular vacuum vessel (40) which contains a plurality of annular superconducting magnets (58). These superconducting magnets are held in a liquid helium holding reservoir (60) such that they are held below their superconducting temperature. A first cold shield (44) and a second cold shield (50 ) have tubular portions between the superconducting magnets and the examination region. These cylindrical portions each include a cylinder (70) of a electrically insulating material such as reinforced plastic. Thermally conductive layers (72) are defined on each surface and are divided by etched slots or resistance portions (74) into a multiplicity of elongated narrow segments (92).

Abstract: An examination region (12) is defined along a z-axis offset from a geometric center (30) of a gradient coil assembly (20). Cylinders of a non-conductive, non-magnetic material support x, y, and z-gradient coils (24, 26, 22) for causing orthogonal magnetic field gradients through the offset examination region. The z-gradient coil (FIG. 2 ) includes a plurality of distributed loop arrays with a winding pattern selected to cause a region of linear magnetic field gradients in the z-direction in a region offset toward a cylinder first end (32). The x and y-gradient coils each include two pairs of oppositely disposed windings (FIG. 3 ) which include a pair of inner spirals (96, 98) offset towards the first end of the cylinder and an outer spiral (90) extending therearound. The outer spiral bows in (92) toward the cylinder first end and fans out (94) toward the second end.

Abstract: An examination region (12) is defined by an elliptically crossed-section cylinder (20) of a non-conductive, non-magnetic material. The x, y, and z-gradient coils (24, 26, 22) are mounted on elliptically crossed-section cylindrical surfaces along the former for causing gradient magnetic fields within the examination region. Each of the x, y, and z-gradient coils includes a plurality of arrays or groups of coil loops. The arrays of coil loops are symmetric about a central plane of symmetry. To either side of the plane of symmetry, each of the first of the x, y, and z coil arrays (74, 80, 90) include at least a first group of loops in which current flows in a clockwise direction and a second array of loops (76, 82, 90) in which current flows in an opposite direction.

Abstract: A planar gradient magnetic field assembly (20) selectively causes gradient magnetic fields linearly across the examination region (12) of a magnetic resonance imaging apparatus. The gradient coil assembly includes a pair of planar y gradient coils (22a, 22b) and a pair of planar x gradient coils (24a, 24b). On each plane, the gradient coil winding includes a pair of larger, outer coil loop arrays (60, 62). A pair of smaller inner coil loop arrays (66, 68) produce a first order gradient field correction to improve the accuracy of the gradient field. The coil loop arrays are symmetric relative to the z axis and relative to an (x,y) plane of symmetry perpendicular to the z axis.

Abstract: A gradient field controller (30) generates current pulses with a preselected profile. Eddy current compensation circuits (42) alter the current pulse profile by adding additional components of selectable frequencies with selectable amplitudes or gains. A power amplifier (32) amplifies the modified current pulse and applies them to gradient field coils (34) of a magnetic resonance imager. A probe or coil (50) monitors the induced gradient response which is integrated (52) to provide an electronic representation of the induced gradient profile. A least squares analysis routine (72) determines the time constant and amplitude of a component attributable to a first eddy current which degrades the induced gradient profile. A filter frequency correction factor calculating routine (76) calculates appropriate filter frequency settings and a gain calibration factor calculating routine (78) calculates the gain settings for the compensation circuits (42).