Koninklijke Philips Corporation N.v.

Marconi Medical Systems, Inc.

Philips Intellectual Property & Standards Gmbh

Steven E Cooke

Thomas Leroy Laurence

Andreas Thon

Frank P DiFilippo

John Vesel

Herfried Karl Wieczorek

Volkmar Schulz

Jerome John Griesmer

Steven J Plummer

Donald R Wellnitz

Carsten Degenhardt

Christopher G Matthews

Jinling Liu

Joseph Lee Molyneux

John F Vesel

A detector array for an imaging system includes an array of scintillator crystals (), wherein each crystal includes a plurality of sides and is laser etched on at least one crystal side to scatter light, and an array of photodetectors () optically coupled to array of scintillator crystals. The side of the crystal () is laser etched with a distinct pattern defined by a user using a computer aided design program. The detector () is part of a nuclear scanner () which includes a reconstruction processor () that reconstructs output signals from the photodetectors () into an image and a user interface () that displays the reconstructed image.

Laser etched scintillation crystals for increased performance

When employing specular reflective material in a scintillator crystal array, light trapping in the crystal due to repetitive internal reflection is mitigated by roughening at least one side () of each of a plurality of pre-formed polished scintillator crystals. A specular reflector material () is applied (deposited, wrapped around, etc.) to the roughened crystals, which are arranged in an array. Each crystal array is coupled to a silicon photodetector () to form a detector array, which can be mounted in a detector for a functional scanner or the like.

Optimized scintilator crystals for PET

When constructing a nuclear detector module in a gantry, a plurality of overlapping light guide modules () are mounted to the gantry in a spaced-apart fashion, and a plurality of underlapping light guide modules () are mounted in between each pair of overlapping light guide modules (). Each of the underlapping modules and the overlapping modules includes a scintillation crystal array () on an interior surface thereof, and a plurality of PMTs on an exterior surface thereof. Overlapping modules () have overlapping structures () that interface with underlapping structures () on the underlapping modules () and thereby eliminate a seam directly beneath PMTs that overlap the crystal arrays of both an overlapping module and an underlapping module. Optical grease is used to form a resilient grease coupling and reduce light scatter between the underlapping and overlapping modules.

PET detector modules utilizing overlapped light guides

An apparatus comprises a plurality of radiation conversion elements () that convert radiation to light, and a reflector layer () disposed around the plurality of radiation conversion elements. The plurality of radiation conversion elements may consist of two radiation conversion elements and the reflector layer is wrapped around the two radiation conversion elements with ends () of the reflector layer tucked between the two radiation conversion elements. The reflector layer () may include a light reflective layer () having reflectance greater than 90% disposed adjacent to the radiation conversion elements when the reflector layer () is disposed around the plurality of radiation conversion elements, and a light barrier layer ().

Radiation conversion elements with reflectors for radiological imaging apparatus

In nuclear imaging, when a gamma ray strikes a scintillator, a burst of visible light is created. That light is detected by a photodetector and processed by downstream electronics. It is desirable to harness as much of the burst of light as possible and get it to the photodetector. In a detector element (), a first reflective layer () partially envelops a scintillation crystal (). The first reflective layer () diffuses the scintillated light. A second reflective layer () and a support component reflective layer () prevent the light from leaving the scintillation crystal () by any route except a light emitting face () of the scintillator (). In another embodiment, a light concentrator () is coupled to the scintillator () and channels the diffuse light onto a light sensitive portion of a photodetector (). The reflective layers () and the concentrator () ensure that all or nearly all of the light emitted by the scintillator () is received by the photodetector ().

Reflector and light collimator arrangement for improved light collection in scintillation detectors

A radiation detector () includes scintillator pixels () that each have a radiation-receiving end, a light-output end, and reflective sides extending therebetween. The reflective sides have a reflection characteristic () varying between the radiation-receiving end and the light-output end such that a lateral spread of light emanating from the light-output ends of the scintillator pixels responsive to a scintillation event generated in one of the scintillator pixels depends upon a depth of the scintillation event in the scintillator pixel. A plurality of light detectors () optically communicate with the light-output ends of the scintillator pixels to receive light produced by scintillation events.

Pixelated detectors with depth of interaction sensitivity

A scintillator () includes radioactive elements which emit radiation of a characteristic energy, such as lutetium, which emits 202 keV and 307 keV &#x3b3;-rays. The scintillators have light output levels that vary and photomultiplier tubes that respond to the light scintillations tend to drift. When a scanner () is not generating diagnostic images, the photomultiplier tubes detect scintillations from the lutetium 176 radiation. A self-calibration processor () adjusts the gain for each photomultiplier tube such that its output peak corresponds to 202 keV or 307 keV and adjusts a scaling factor for PMT outputs corresponding to each scintillator such that the output peaks have a common amplitude.

PMT gain and energy calibrations using lutetium background radiation

A method of reconstructing time-of-flight (TOF) images includes obtaining a profile of a subject to be imaged in an examination region () of an imaging system (), Events associated with radiation emitted from the subject are detected and converted to electronic data. Electronic data attributable to radiation events located outside the profile are removed and images are reconstructed from the remaining electronic data.

Method of constructing time-in-flight pet images

A method of locating an event with a gamma camera () of an emission computed tomography (ECT) scanner () is provided. The gamma camera () includes a matrix of sensors () situated to view the event. The sensors () have respective outputs that are responsive to the event. The method includes: identifying a first sensor in the matrix that has in response to the event a highest output relative to the other sensors in the matrix (step (B)); identifying a number of second sensors in the matrix that are closest neighbors to the first sensor (step (B)); combining into a total output a number of outputs from the identified sensors, the number of outputs being at least one (1) and less than the number of all the identified sensors (step (B)); and, determining a threshold value which is a percentage of the total output (step (B)).

Gamma camera with dynamic threshold

A nuclear camera system includes a detector ( ) for receiving radiation from a subject ( ) in an exam region ( ). The detector ( ) includes a scintillation crystal ( ) that converts radiation events into flashes of light. An array of sensors ( ) is arranged to receive the light flashes from the scintillation crystal ( ). Each of the photomultiplier sensors ( ) generates a respective sensor output value in response to each received light flash. A processor ( ) determines when each of the radiation events is detected. At least one of an initial position and an energy of each of the detected radiation events is determined in accordance with respective distances (d . . . d ) from a position of the detected event to the sensors ( ). An image representation is generated from the initial positions and energies.

Event localization and fall-off correction by distance-dependent weighting

In positron emission imaging, coincident gamma ray pairs are acquired and processed to generate an image. Random gamma ray pairs in the acquired coincidence data degrade the quality of the resultant image. The coincident gamma ray pairs are re-paired to generate non-coincident gamma ray pairs. The non-coincident pairs are used to correct for randoms in the acquired coincidence data. Alternately, singles gamma rays may be detected and paired with non-coincident single gamma rays to generate non-coincident pairs. These pairs may be used to correct for randoms in the acquired coincidence data.

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Garfield Heights, OH, United States of America

Steven E Cooke