Radiant energy – Invisible radiant energy responsive electric signalling – With or including a luminophor
Reexamination Certificate
1999-09-17
2002-08-06
Porta, David P. (Department: 2882)
Radiant energy
Invisible radiant energy responsive electric signalling
With or including a luminophor
C250S370090
Reexamination Certificate
active
06429434
ABSTRACT:
TECHNICAL FIELD
This invention relates to the field of tomography. More specifically, the present invention relates to a method of measuring and correcting the attenuation associated with detecting coincidences using a collimated source and a dedicated detector for improved measurement sensitivity.
BACKGROUND ART
Positron Emission Tomography (PET) has gained significant popularity in nuclear medicine because of the ability to non-invasively study physiological processes within the body. Applications employing the PET technology for its sensitivity and accuracy include those in the fields of oncology, cardiology and neurology.
Using compounds such as
11
C-labeled glucose,
18
F-labeled glucose,
13
N-labeled ammonia and
15
O-labeled water, PET can be used to study such physiological phenomena as blood flow, tissue viability, and in vivo brain neuron activity. Positrons emitted by these neutron deficient compounds interact with free electrons in the body area of interest, resulting in the annihilation of the positron. This annihilation yields the simultaneous emission of a pair of photons (gamma rays) approximately 180° (angular) apart. A compound having the desired physiological effect is administered to the patient, and the radiation resulting from annihilation is detected by a PET tomograph. After acquiring these annihilation “event pairs” for a period of time, the isotope distribution in a cross section of the body can be reconstructed.
PET data acquisition occurs by detection of both photons emitted from the annihilation of the positron in a coincidence scheme. Due to the approximate 180° angle of departure from the annihilation site, the location of the two detectors registering the “event” define a chord passing through the location of the annihilation. By histogramming these lines of response (the chords), a “sinogram” is produced that may be used by a process of back-projection to produce a three dimensional image of the activity. Detection of these lines of activity is performed by a coincidence detection scheme. A valid event line is registered if both photons of an annihilation are detected within a coincidence window of time. Coincidence detection methods ensure (disregarding other second-order effects) that an event line is histogrammed only if both photons originate from the same positron annihilation.
In the traditional (2-D) acquisition of a modern PET tomograph, a collimator (usually tungsten) known as a septa is placed between the object within the field-of-view and the discrete axial rings of detectors. This septa limits the axial angle at which a gamma ray can impinge on a detector, typically limiting the number of axial rings of detectors that a given detector in a specific ring can form a coincidence with to a few rings toward the front of the tomograph from the given detector's ring, the same ring that the detector is within, and a few rings toward the rear of the tomograph from the given detector's ring.
The current state of the art nuclear medicine camera is capable of operating in either PET or Single Photon Emission Computed Tomography, (“SPECT”), modes. Fundamentally, in both modes, gamma photons emitted from within a patient are detected. However, a significant portion of these emitted photons are obstructed from reaching the detectors by colliding with atoms. When this occurs, one significant possibility is a course alteration away from the detector that may result in a missed detection. The degree of attenuation depends upon the amount and density of matter between the emitting source and the detector, and will vary from subject to subject. The more attenuation present, the less probable will be the accurate detection of a gamma photon. Unless the amount of attenuation is known, the detected activity within a defined energy window underestimates the true activity. This results in poorer contrast and attenuation artifacts in the reconstructed images. Conditions such as these reduce the confidence one may have in extracting information for diagnosis. However, attenuation methods well suited for SPECT are poorly suited for PET.
To compensate for this phenomenon, it is now common to incorporate an apparatus to transmit gamma photons of a known flux density through a patient so that the patient induced attenuation can be measured. Attenuation was first measured in PET by using a ring of positron emitting isotope surrounding the object to be measured. In this technique, the ratio between a transmission scan and a blank scan form the attenuation. The blank is measured by simply measuring the rate that gamma rays from positrons are detected by the detection system when no attenuating media is present. In the original scanners as described above as having septa, the septa are provided for collimating the gamma rays in an axial direction, but the rings allow for no transaxial collimation. The lack of collimation allow the acceptance of scattered events into the transmission measurement, resulting in an underestimate of the attenuation. To improve the transmission measurement, systems use rotating rod sources. These sources are disposed in parallel fashion to the axis of the scanner and are collimated in the axial direction by the septa. In the transaxial direction, the collimation may be provided electronically since the position of the source is known. However, the activity in the rod must be the same as that activity in the earlier ring source to provide the same count rate. With modern block detectors, the dead-time of the near block limits the activity in the rod.
A more recent advancement in PET acquisition is 3-D, in which the septa are removed, which allows a given detector to be in coincidence with detectors from all other detector rings. With the advent of three-dimensional reconstruction techniques, greater sensitivity to emission counts is possible if the septa are removed. As the septa represent a significant cost, there is also an economic incentive to exclude them from the system. However, with the absence of septa, the problems of both detector dead-time and scatter are magnified.
Since the position of a source with respect to the detector system can be known, there is no need to detect coincidences, thereby allowing the use of a source that emits single gamma rays. Only one detector—the detector on the far side of the system—is needed to make the transmission or blank measurements. Without the counting losses due to the dead-time of the near detector, the activity of the source may be increased resulting in an increase in count-rate and thus a better quality measurement. However, without axial collimation, the scatter included in the transmission scan causes an underestimate of the attenuation measurement. To decrease the possibility of scatter, the gamma rays from the source can be collimated with lead or tungsten to form a beam that illuminates only a narrow plane of detectors. Other gamma rays that would only contribute to background are eliminated. Since the directionality of single gamma rays cannot be determined, only a single point of activity illuminating a detector bank can be used. This requires increased levels of activity to meet the count-rate needed for an adequate quality measurement. Also, the scanning protocol is more efficient if the transmission measurement is performed after the patient has been injected with radioactivity. Even though a different isotope such as
137
Cs which emits gamma rays with an energy of 662 keV can be used for the transmission scan, there is a significant difficulty in distinguishing the transmission events from the emission events.
(SPECT) is similar to PET. However, in SPECT, only a single photon from a nuclear decay within the patient is detected. Also, the line of response traveled by the photon is determined exclusively by detector collimation in SPECT, as opposed to the coincident detection of two collinear photons as in PET.
There are a number of correction methods in the art. However, these methods utilize SPECT based approaches for transmission based attenuation correction th
Casey Michael E.
Hamill James J.
Miller Stephen D.
Nutt Ronald
Watson Charles C.
Hobden Pamela R.
Pitts & Brittian P.C.
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