SPECT for breast cancer detection

Details SPECT for breast cancer detection SPECT for breast cancer detection

2003-04-30

2004-09-21

Epps, Georgia (Department: 2873)

Radiant energy

Invisible radiant energy responsive electric signalling

With or including a luminophor

C250S370090

Reexamination Certificate

active

06794653

ABSTRACT:

FIELD OF THE INVENTION
The present invention relates generally to SPECT instruments and in particular to a small SPECT device, dedicated to breast imaging.
BACKGROUND OF THE INVENTION
SPECT is one of several nuclear imaging techniques. Generally, in nuclear imaging, a radioactive isotope is injected to, inhaled by or ingested by a patient. The isotope, provided as a radioactive-labeled pharmaceutical (radio-pharmaceutical) is chosen based on bio-kinetic properties that cause preferential uptake by different tissues. The gamma photons emitted by the radio-pharmaceutical are detected by radiation detectors outside the body, giving its spatial and uptake distribution within the body, with little trauma to the patient.
FIG. 1
illustrates a general nuclear-imaging detector
10
comprising a NaI(T1) scintillation crystal
12
. Generally, scintillation crystal
12
, of a diameter D
1
, is large enough to image a significant part of the human body (typically 40 cm). An array of photo-multiplier tubes (PMTs)
14
view scintillation crystal
12
, to give positional sensitivity. Each PMT
14
has an x and a y coordinate. When a photon is absorbed by scintillation crystal
12
, light is generated. A number of PMTs
14
receive the light and produce signals. The X and Y coordinates of the event are determined by the strength of the signals generated by each PMT. The energy of the event is proportional to the sum of the signals, called the Z signal. Only Z signals within a given range are counted.
Semiconductors with high atomic numbers and relatively high densities such as CdZnTe, CdTe, HgI
2
, InSb, Ge, GaAs, Si, PbCs, PbCs, PbS, or GaAlAs, have a high stopping power and can be used as gamma ray detectors with good photon detection efficiencies, good spatial resolution, and a relatively high photon-energy resolution. Solid state semiconductor gamma cameras generally comprise arrays of pixelated detector, hereinafter referred to as “pixelated detectors”. One type of pixelated detector is described in PCT publication WO 98/23974, the disclosure of which is incorporated herein by reference.
FIG. 2
shows a typical construction of a pixelated detector
20
comprising a crystal
22
formed from a semiconductor material such as one of those noted above. A face
24
of crystal
22
has a large single cathode electrode
26
. An opposite face
28
of crystal
22
has an anode
30
comprising a rectangular array of identical small square anode pixels
32
. Typically, sizes of anode pixels
32
vary between 1 and 4 mm
2
, and the thickness of crystal
22
, between anode
30
and cathode
26
is on the order of millimeters to a centimeter. In operation, a voltage difference is applied between anode and cathode so that an electric field, hereinafter referred to as a “detector field”, is generated in crystal
22
. This field is typically on the order of a few kilovolts per centimeter.
When a photon, having an energy typical of the energies of photons used in gamma cameras, is incident on crystal
22
, it generally loses all its energy in crystal
22
by ionization and produces pairs of mobile electrons and holes in a localized region of crystal
22
. As a result of the detector field, the holes drift to cathode
26
and the electrons drift to anode
30
, thereby inducing charges on anode pixels
32
and cathode
26
. The induced charges on anode pixels
32
are sensed and generally partially processed by appropriate electronic circuits located in a detector base
34
to which detector
20
is mounted. Signals from the induced charges on pixels
32
are used to determine the time at which a photon is detected, how much energy the detected photon deposited in the crystal and where in the crystal the photon interaction took place.
Generally, a collimator
16
is placed between scintillation crystal
12
or
22
and the tissue. Commonly, collimator
16
is honeycomb shaped, comprising a large number of holes separated by parallel lead septa. The purpose of collimator
16
is to intercept and eliminate gamma-ray photons that are not traveling in an accepted direction, parallel to the lead septa. The geometric configuration of collimator
16
(e.g., the size and shape of the holes and the length of lead walls) determine the geometric response function of collimator
16
. In general, there is a tradeoff in which the collection efficiency of detector
16
increases as the geometric response function widens and the spatial resolution decreases.
SPECT (Single-Photon-Emission Computed Tomography) is based on conventional nuclear imaging technique and tomographic reconstruction methods, wherein projection (or planar) data acquired from different views around the patient are reconstructed, using image reconstruction methods, to generate cross-sectional images of the internally distributed radio-pharmaceuticals. SPECT images provide enhanced contrast, when compared with planer images obtained with conventional nuclear imaging methods.
A typical SPECT system consists of a single or multiple units of radiation detectors arranged in a specific geometric configuration, a mechanism for moving the radiation detectors and/or a specially designed collimator to acquire data from different projection views.
A typical system is based on a single or multiple scintillation cameras, mounted on a rotating gantry. This may consist of a single-camera system, a dual-camera system, a triple-camera system or a quadruple-camera system. Generally, camera-based SPECT systems use Anger scintillation cameras such as those used in conventional planar nuclear medicine. These cameras allow for truly three-dimensional imaging, by providing a large set of contiguous trans-axial images that cover the entire organ of interest. They are easily adaptable for SPECT imaging of the brain or body, by simply changing the radius of rotation of the camera. Because of the relatively low counting rate capability, it is desirous to use two or more cameras. The increase in sensitivity per slice is proportional to the number of cameras.
The most common collimator systems for SPECT systems are parallel-hole collimators, wherein the septa are perpendicular to the scintillation crystal. As noted above, there is an unavoidable tradeoff between detection efficiency and spatial resolution of parallel-hole collimators. An alternative design is a converging-hole collimator that increases the angle of acceptance of incoming photons. Examples of these special collimator designs and their points of convergence are given in
FIGS. 3A-3D
.
FIG. 3A
illustrates a parallel-hole collimator.
FIG. 3B
illustrates a fan-beam collimator, where the collimator holes are converged to a line that is parallel to the axis of rotation.
FIG. 3C
illustrates a cone-beam collimator, where the collimator holes are converged to a point.
FIG. 3D
illustrates a varifocal collimator, where the collimator holes are converged to various focal points. Converged collimators have smaller fields of view when compared with parallel-hole collimators.
In general, camera-based SPECT systems may be comprised of detectors that rotate about the region to be imaged or arrays of detectors that completely surround the region to be imaged, for example with a ring of detectors. In either event, the object of the imaging system is to acquire data from directions including at least 180 degrees of view.
SPECT breast imaging is especially challenging, since there is no way to acquire cross-sectional data suitable for analytic tomographic reconstruction, especially from slices near the chest wall, from the front of the patient, since there is no way to acquire views of a slice over 180 degrees of view. Furthermore, rotating around the axis of the patient is not desirable due to the distance of the back of the body from the breast and the amount of radiation from the body which may overwhelm the radiation emitted from the breast.
SUMMARY OF THE INVENTION
An aspect of some embodiments of the present invention relates to providing a small SPECT system, dedicated to the nuclear imaging of the breast. In exemplary e

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