Abstract: A CT scanner includes a stationary gantry (C) having an examination region (12) centrally therein. A rotating frame or gantry (C) which is mounted for rotation about the examination region carries an x-ray tube assembly (B), a liquid-to-liquid heat exchanger (34), and a liquid-to-air heat exchanger (48) around a peripheral edge thereof. A first closed loop (30) carries a first cooling fluid, particularly oil, between a housing (22) which surrounds an x-ray tube (68) and the liquid-to-liquid heat exchanger to remove heat from the x-ray tube. A second closed loop (40) conveys a second cooling fluid, particularly water, between the liquid-to-liquid heat exchanger and the peripheral liquid-to-air heat exchanger. The second closed loop includes a reservoir (44) for storing a substantial volume of water such that a significant portion of the heat generated during an x-ray examination can be stored by the water in the reservoir.

Abstract: Granulated bismuth is mixed with a liquid carrier into a suspension. The carrier - bismuth combination is applied to a surface in a thickness sufficient to provide a desired amount of radiation attenuation. The carrier adheres to the surface and maintains the bismuth adjacent said surface.

Abstract: In a magnetic resonance imaging system, a read gradient scaler (102) scales the amplitude and width of read gradients (78, 82, 84). A sampling rate control (104) controls the sampling rate of the resonance signals received from corresponding magnetic resonance echoes (80 54, 86). For example, when the amplitude of the gradient pulse is doubled and its width halved, the sampling rate of the resultant magnetic resonance signal is doubled, e.g., from a bandwidth of 32 MHz to a bandwidth of 64 MHz. In this manner, some echoes are read-out over a longer period of time with a lower bandwidth to produce lower signal-to-noise data lines; whereas, other echoes are much shorter and are read-out more quickly, but with a lower signal-to-noise ratio.

Abstract: An x-ray tube of a CT scanner is powered by a high-voltage power supply (26). The high-voltage power supply includes a plurality of sections (102) each having three straight-up transformers (48) which receive three 120.degree. phase shifted alternating current components as inputs. The straight-up transformers perform a direct voltage transformation with single or multiple transformers and with no capacitive multipliers. Each straight-up transformer has a primary winding (T1) and two secondary windings (T1-A, T1-B). The secondary windings are connected together in delta and wye configurations (84). The alternating current components have their voltage boosted and are rectified and summed to form a high-voltage output that is substantially ripple-free. A pulse-width modulated converter (34) generates a conditioned output current from an inputted direct current.

Abstract: In a magnetic resonance imaging apparatus, a radio frequency coil (40) is disposed closely adjacent the patient's head. The radio frequency coil includes a first annular ring (80, 114) around the patient's head from which a first plurality of legs (82, 116) extend. Opposite legs are interconnected equidistant from the first annular ring to form a virtual ground connection (84, 118). In the embodiment of FIG. 6, a second annular ring (120) is disposed parallel to the first annular ring with a second plurality of legs (122) extending between the first and second annular rings. By adjusting a ratio .rho. of the current flow in the loops defined by the first legs, the first annular ring and virtual ground relative to the current loops defined by the second legs and the first and second annular rings, the linearity of the B.sub.1 field within the head coil is selectively adjustable (FIG. 8).

Abstract: A SPECT system includes two radiation detector heads (32) and (34) which are mounted opposite each other to a gantry (30) for rotation about a subject. The subject is injected with a source of emission radiation, which emission radiation passes through a collimator (38) mounted on each detector head and is received by a radiation receiving face of the detector heads. A radiation source (40) is disposed between the collimator (38) of one detector head and the face of the one detector head. Transmission radiation from the radiation source (40) passes through the collimator (38), through the subject, and is received by the opposite detector head concurrently with the emission radiation. Alternately, a radiation source (40) is disposed behind the collimator of both detector heads (32, 34) such that both detector heads concurrently receive emission and transmission radiation.

Abstract: A resonance exciting coil (C) disposed in an image region in which a main magnetic field and transverse gradients have been produced. A flexible receiving coil (D) includes a flexible plastic sheet (40) on which an electrically continuous flexible foil strip (36) is adhered to receive signals from the resonating nuclei. The coil also includes components (24), mounted on the flexible plastic sheet (40), which may amplify the received signals before they are transmitted along a cable (22). A first soft material layer (44) is mounted on the flexible plastic sheet (40) for providing comfort to the patient. A flexible mechanical structure (50) on which two flexible receiving coils (F, G) adjusts an overlap between the coils (F, G) as the coils are flexed to adjust interaction properties of the coils (FIGS. 4 A, 4B).

Abstract: Magnetic resonance is excited (50) in first and second species dipoles of a subject in a temporally constant magnetic field. The resonance is refocused (52) to generate a spin echo (54) centered at a time when the first and second species resonance signals are in-phase. Gradients echoes (64, 68) are generated, centered at a time (2n+1).pi./.delta..omega. before and after the spin echo, where .delta..omega. is a difference between the first and second species resonance frequencies. In this manner, the first and second species signals are 180.degree. out-of-phase in the gradient echoes. The resonance is refocused (82) one or more times to generate additional spin and gradient echoes with different phase encodings (78). The sequence is repeated with yet more phase encodings, and magnetic resonance signals from the spin echo and the two gradient echoes are reconstructed (86) into a spin echo image (s.sub.0) and a pair of gradient echo images (s.sub.+1, s.sub.-1).

Abstract: A magnetic resonance imaging apparatus includes a magnet assembly (10) for generating a temporally constant primary magnetic field through a central bore (14). The central bore is surrounded by a gradient coil assembly (30) including a dielectric former (32) and gradient coils (34, 36, 38) for generating magnetic field gradient pulses across the examination region. A radio frequency coil assembly (50) including a birdcage coil (52) is mounted inside of the gradient coil assembly and surrounding the examination region. A radio frequency shield (60) includes a dielectric sheet (62) having a plurality of first metal foil strips (72) defined on each surface. The metal strips are defined by etching or cutting gaps (70) in a continuous sheet of copper foil, leaving bridges (74) across the gaps adjacent opposite ends of the foil strips.

Abstract: A CT scanner (10) includes a reconstruction processor (82) for reconstructing an image from digital signals from detector arrays (20). Each detector array includes scintillation crystals (22) arranged in an array for converting x-ray radiation into visible light. An array of photodetectors (24) is mounted beneath the scintillation crystal array for converting light emitted from the scintillation crystals into electrical charge. The combination scintillation crystal, photodetector array is mounted to and preferably integral with an integrated circuit (26) having charge storage devices (32, 34). Electrical charge generated by each photodetector is integrated and stored alternately by a corresponding pair of charge storage devices. The charge storage devices operate as a double buffer in which one charge storage device accumulates charge while the other holds an accumulated charge until read out by downstream circuitry.

Abstract: A medical diagnostic imaging system includes a system controller (16) which generates first logic format control signals. The first logic format control signals operate an x-ray source and are designed to operate a large format camera through a plurality of data paths. The large format camera is disconnected from the diagnostic imaging system and replaced with digital camera (32) and a digital imaging processor (34) which are incompatible with the diagnostic imaging system. A universal interface (36) is selectively connected to the data paths of the diagnostic imaging system and to the digital imaging processor (34). The universal interface (36) intercepts the first control signals from the system controller (16), and generates second control signals based on the first control signals which operate the digital imaging processor (34) and digital camera (32) transparent to the operations of the system controller (16).

Abstract: A patient support (40) has lower sections (50a, 50b) of first and second gradient coil assembly portions (42a, 42b) affixed thereto. Upper sections (52a, 52b) of the gradient coil assembly portions are selectively removable from the lower gradient coil assembly sections. The gradient coil assembly sections are mounted such that an interstitial gap is defined therebetween. The gradient coil assembly portions are configured to fit snugly around the patient's torso below the shoulders and around the patient's head, but are too small in diameter to pass around the patient's shoulders. The radio frequency coil (44) has a lower portion (60) disposed below the patient's shoulders and an upper portion (62) removably positionable above the patient's shoulders. The gap between the two gradient coil assembly portions is preferably between 1/2 and 3/4 of a diameter of the gradient coil portions.

Abstract: A magnetic resonance imaging apparatus includes main field coils (10) for generating a temporally uniform magnetic field longitudinally through a central bore (12). A whole body gradient magnetic field coil (30) and radio frequency coil (36) are disposed around the bore. An insertable coil assembly (40) includes an insertable gradient coil (42) and a radio frequency coil (44). The insertable coil assembly place the gradient and radio frequency coil significantly closer to the region of interest to be imaged than the whole body coils. In the embodiment of FIG. 2, the gradient coil assembly is a cylindrical coil assembly including x, y, and z-gradient coils which receive current pulses in order to generate linear, uniform, magnetic field gradients through a selected region adjacent a patient end of the gradient coil.

Abstract: An x-ray tube includes an anode (A) and an envelope (C). A cathode assembly (B) which is supported in the envelope on a bearing (32) emits a beam of electrons which strike the anode forming a focal spot. The anode rotates (D) relative to the cathode such that focal spot follows a generally annular path along a beveled track (14). If the axis of the anode and the cathode assembly are screwed or offset, the focal spot path is not circular and wobbles. An adjustment assembly (60) adjusts the relative positions of the anode, the cathode and the envelope to adjust the anode and cathode assembly axes. The adjustment assembly also includes one or more electrodes (102, 108) which adjust the position of the focal spot. An angular position encoder (106) identifies an angular orientation of the anode. A control circuit (110) applies an electrostatic potential to the electrodes to move the focal spot such that it stays on a constant plane of the leveled anode surface.

Abstract: A vertical B.sub.0 temporally constant magnetic field is defined between a pair of pole faces (12, 14) that are interconnected by a C-shaped ferrous magnetic flux path (16). A quadrature radio frequency coil array (50) is disposed in a plane orthogonal to the B.sub.0 field. The coil array includes a plurality of coils (50.sub.1, 50.sub.2, . . . ) that are disposed in a partially overlapping relationship. Each of the coils has a peripheral loop (60), preferably defined by four linear legs (60.sub.1, 60.sub.2, 60.sub.3, 60.sub.4) of equal length which define a square. A pair of crossing elements (62.sub.1, 62.sub.2) are connected with mid-points of opposite sides of the square, the opposite mid-points are 180.degree. out-of-phase with each other at the magnetic resonance frequency and 90.degree. out-of-phase with neighboring mid-points of the square. The crossing elements cross but are not connected, in a symmetric relationship. Each of the crossing elements has a radio frequency pick-up (64.sub.1, 64.sub.

Abstract: An x-ray beam limiter includes a compressible disc 18 containing an aperture 20 and a slot 22 extending from the aperture to the outer edge 28 of the disc. The outer edge 28 of the disc 18 is covered with a bearing material and is retained within a groove 36 in a bearing retaining race 32 such that the disc 18 is rotatable within the retaining race 32. A beam limiter 54 is mounted to the disc 18. The disc 18 and beam limiter 54 are rotatable between at least two different positions.

Abstract: A SPECT camera system has two detector heads (16a, 16b) which are movable between position (FIG. 2) and an adjacent, orthogonal position (FIG. 3). The detector heads have slide members (42a, 42b) which are slidably received on linear guide members (40a, 40b) of carriers (16a, 16b). The carriers have followers (20a, 22a, 24a; 20b, 22b, 24b) which are slidably received along linear guide paths (26a, 28a, 30a; 26b; 28b, 30b). Hydraulic or other extensible cylinders (32a, 32b) selectively permit the carriers, hence the detector heads, to slide along the guide paths between the 180.degree. opposite and the adjacent, orthogonal positions. Preferably, a rotating gantry (12) is rotated to place both detector heads above an intended position. In this manner, the extensible cylinders controllably lower the detector heads as they move by gravity from the present position to the intended position. In both the 180.degree.

Abstract: A movable patient supporting portion (10) of a patient couch (A) includes a socket (26) for receiving a mating plug (24) on a localized coil (B). The patient couch selectively inserts the localized coil and a supported patient into a bore (14) of a cryogenic magnet system (C). The localized coil includes a resistor (86) whose magnitude identifies the coil. A coil identification interrogator (84) interrogates the coil identification resistor and derives a corresponding binary coil identification. The coil identification addresses a look-up table (90) to retrieve diagnostic test information, an identification of a coil for a human-readable display, and, preferably, an identification of an isocenter of the coil. A diagnostic test unit (92) electrically tests the coil through the plug and socket connection with the diagnostic tests prescribed by the look-up table.

Abstract: A SPECT camera system has detector heads (14a, 14b, 14c) each having a collimator (18) facing toward an examination region (10). The detectors receive emission radiation from a radioisotope injected into a subject in the examination region and, preferably, also receive transmission radiation from a transmission source (20) disposed across the examination region. The detectors rotate around the subject through a plurality of projection angles. The detectors generate projection views from the received radiation for each projection angle. The projection views are stored in a projection view memory (26). A reconstruction processor (38) reconstructs the projection views into a volumetric image representation that is stored in a volumetric image memory (40). A reprojector (42) reprojects the volumetric image and volume of interest along each projection angle producing a set of reprojection views and regions of interest.

Abstract: To calibrate a nuclear camera, the scintillation crystal (10) is irradiated with a uniform flood source (34). Scintillation events cause corresponding electrical pulses from photomultiplier tubes (12) that are optically coupled to the scintillation crystal. A selection circuit (40) compares (42) the amplitude or energy of pulses at each coordinate with a median energy and divides the pulses to generate a higher energy image (44) and a lower energy image (46). When the higher and lower energy images are subtracted (50) regions of the difference image (52) with high count densities identify a photomultiplier tube for gain adjustment. A gain adjustment circuit (60) monitors the amplitude of the electrical pulses and the number of pulses with each amplitude from the selected photomultiplier tube to generate a raw data energy distribution curve (FIG. 3). The energy distribution curve is smoothed (70) and the first derivative is taken (72). The curve is filtered (74) to isolate a preselected energy region.