Koninklijke Philips Corporation N.V.

Adac Laboratories, Inc.

Ugm Laboratory, Inc.

University of Pennsylvania

Siemens Gammasonics, Inc.

Ugm Medical Systems, Inc.

Koninklijke Philips Electronics N.v.

Gerd Muehllehner

Michael J Geagan

Matthew J Murphy

Horace H Hines

Walter Ruetten

Andreas Thon

Donald R Wellnitz

Everett W Stoub

Klaus Fiedler

Peter Nellemann

Edward M Goldberg

Joel Karp

William K Braymer

Robert M Lewitt

James G Colsher

Michael J Parma

Peter Countryman

Joel S Karp

A time-of-flight PET nuclear imaging device (A) includes radiation detectors (), electronic circuits () for processing output signals from each of detectors (), a coincidence detector (), a time-of-flight calculator () and image processing circuitry (). A calibration system () includes an energy source () which generates an electrical or optical calibration pulse. The electrical calibration pulse is applied at an input to the electronics at an output of the detector and the optical calibration pulse is applied to a preselected point adjacent a face of each optical sensor () of the detectors. A calibration processor () measures the time differences between the generation of the calibration pulse and the receipt of a trigger signal from the electronic circuitry by the coincidence detector () and adjusts adjustable delay circuits () to minimize these time differences.

Timing calibration for TOF-PET scanner

A cylinder () of radiation detectors () in a diagnostic imaging apparatus detects radiation pairs with corresponding electronic detector channels () which have non-uniform time-varying temporal delays. A plurality of calibration radiation sources () that concurrently emit radiation pairs is mounted at known positions inside of the cylinder. A temporal correction processor () uses the known positions of the calibration sources to determine errors in the relative detection times of radiation pairs from the calibration sources and uses the determined errors to generate corrections for the relative detection times of radiation pairs from radiopharmaceuticals injected in an imaged subject.

Timing calibration using radioactive source

A radiographic imaging system () includes an array of detectors () arranged around a circular bore (). A subject () is injected with a radioisotope. The subject (), supported by a subject support means (), is to be placed into the bore () for an examination. A fixed, non-circular shield () is rigidly mounted to one of an entrance () and an exit () of the bore () to prevent emission radiation originating outside of the bore () to reach the radiation detectors (). The shield () extends from an outer periphery of the bore () toward and surrounding a central axis of the bore () and defines a fixed, non-circular subject receiving aperture (). The shield () is tailored to a contour of the subject support and the subject () to permit the subject of a maximum girth to be received in the fixed, non-circular aperture ().

Image system with non-circular patient aperture

In positron emission tomography, a nuclear medicine scanner is utilized to detect &#x3b3;-ray events resulting from positron annihilation events. Molecules with known behaviors are tagged with radioactive isotopes which decay into &#x3b3;-ray pairs which are detected coincidentally, i.e. in a near-simultaneous fashion, by radiation detectors. A temporal recorder and a subject support monitor indicate the time and position of the subject when the coincident &#x3b3;-rays were detected. A storage buffer collects &#x3b3;-ray detection times and locations along with support positions. Every 1/100- 1/10second, a batch of data collected in the buffer is reconstructed into overlapping portions of an image memory as the support moves continuously through the scanner.

Real-time list mode reconstruction

A positron emission detection scanner includes a first plurality of detecting elements arranged in a first two dimensional geometrical array, the detecting elements defining a first detection surface oriented for receiving radiant energy stimulus incident thereto. The detecting elements each have a second surface for communicating light from a scintillation event occurring in response to receiving a radiant energy stimulus. A light transmitting member is provided for receiving light from the scintillation events from each of the detecting elements. A second plurality of light sensing members is arranged in a second two dimensional geometrical array, different from the first geometrical array, the alignment of the light sensing members is independent of the detecting elements. A predetermined group of the light sensing members is responsive to light from said light transmitting member with each one of the light sensing members of the group producing signals proportional to its respective portion of the collected light.

Pixelated scintillation detector

In the present invention, a tomographic emission scanner has a detection chamber with an axial center line and an axial periphery, and a plurality of single element emission detector units are positioned around the axial periphery of the detection chamber. Each detector unit includes a single detecting element having a detection face oriented toward the detection chamber. The detection face of each detecting element is curved along the axial periphery of the detection chamber.

Tomographic emission scanner having curved element detectors

A tomographic emission scanner has a detection chamber with an axial center line and an axial periphery, a positron emission coincidence detection circuit, and a number of positron emission detector units that are positioned around the axial periphery of the detection chamber. Each detector unit is operatively coupled to the detection circuit and has a detection face oriented toward the detection chamber. Each detector unit is circumferentially movable about the axial periphery of the detection chamber. In one embodiment, at least one of the detector units is independently movable with respect to any other detector unit. In another embodiment, the scanner also has a resting surface at least partially positioned within the detection chamber. The resting surface extends parallel to and is radially movable with respect to the axial center line. The resting surface is also circumferentially movable with respect to the axial center line while continuously facing upward.

Tomographic emission scanner having offset high detection efficiency area

A method of reconstructing an image in a medical imaging system comprises the steps of detecting transverse rays and oblique rays using a limited range of projection angles. The limited range of projection angles is achieved by either limiting the range of rotation of the detectors at each position along the z axis, or by causing the detectors to trace a continuous-motion helical path along the z axis. Data is then rebinned to create a stack of partially-complete sinograms corresponding to individual transverse slice images corresponding to, in the aggregate, a three-dimensional image. Because of the limited projection angles, each sinogram represents an incomplete data set. The rebinning includes, for each oblique ray, identifying among the sinograms the images intersected by the ray, and, applying to each of the sinograms associated with the images intersected by the ray an increment that is representative of the ray. A two-dimensional iterative reconstruction algorithm is applied to each partial sinogram in order to reconstruct a three-dimensional image slice-by-slice, followed or preceded by axially filtering to reduce blurring in the z direction.

Image generation from limited projections in positron emission

A gamma camera for a Positron Emission Tomography (PET) system comprises a single, continvous scintillation crystal having an imaging surface, and a first layer of septa disposed along the imaging surface between the object and the imaging surface. A gap is provided between the imaging surface and the first layer of septa to allow non stray radiation to reach the imaging surface. The septa may be arranged to have a long axis disposed at an acute angle away from perpendicular to an axis of rotation or made to move in relation to the imaging surface of the scintillation crystal to reduce cold spots on the image. Multiple layers of septa may be provided.

Gamma camera having partial septas and moving septas for positron

A system including circuitry within a gamma camera system for generating spatially variant photomultiplier cluster based on a peak photomultiplier and for generating spatially variant weights for photomultipliers of the cluster depending on a cluster type signal. The system includes a first memory circuit addressed by a peak photomultiplier address signal and responsive thereto which generates (1) a unique photomultiplier cluster in geometry and size for the peak photomultiplier and (2) a cluster type. A resolution input signal (high/low) changes the size of the cluster. The first memory is programmable. A second memory responsive to photomultiplier addresses of the cluster and the cluster type, generates weight values for each photomultiplier. The second memory is programmable. The second memory allows weights for individual photomultipliers to be altered based on the geometry and size (e.g., type) of the cluster generated by the first memory. The system can effectively operate within a digital gamma camera system.

Spatially variant PMT cluster constitution and spatially variant PMT

Images are reconstructed using multi-slice rebinning of raw data to create a stack of two-dimensional data sets. A three-dimensional image is reconstructed slice-by-slice from the data sets and the rebinned data is axially filtered to reduce blurring resulting from the rebinning, the filtering being performed either before or after reconstruction of the three-dimensional image.

Three dimensional image reconstruction for a positron emission tomograph

A method and apparatus for correcting for the spatial distortions of scintillation cameras or similar image forming apparatus. The spatial distortion correction method accurately and precisely determines distortion correction factors in an off-line test measurement and analysis phase prior to actual on-line diagnostic use. The distortion correction factors are initially determined from image event data that is obtained during the test measurement phase by orthogonal line pattern images. Data from a uniform field flood image is also utilized during the test measurement phase to refine and modify the distortion correction factors utilizing the gradient of a function of the field flood image data. Scintillation camera image repositioning apparatus is provided and includes a memory having the correction factors stored therein in a predetermined array. During the on-line use of the scintillation camera and associated image display system, the image repositioning apparatus transforms the data from the scintillation camera in accordance with the stored distortion correction factors. In a preferred arrangement, the image event data is repositioned or corrected event by event.

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Wayne, PA, United States of America

Gerd Muehllehner