Patent Description:
This disclosure relates generally to medical imaging, and more specifically to a radiation detector for a medical imaging scanner.

<CIT> discloses a detector element for spatially resolved detection of gamma radiation. The detector consists of two or more different conversion units which react to the absorption of a gamma quantum with light emissions of different spectral composition. A photodetector arrangement may therefore discriminate between the sites of origin of the light emissions by means of their spectral characteristics. <CIT> discloses a radiation detection apparatus comprising a sensor panel including a sensor unit disposed on a plurality of photoelectric converters on a substrate, a first scintillator layer disposed on the sensor panel, and a second scintillator layer disposed on the first scintillator layer. Further, <CIT> discloses an energy sensitive detector comprising an array of detector elements, wherein the detector elements comprise a first scintillator configured to emit photons within a first wavelength range when stimulated by X-rays, a second scintillator configured to emit photons within a second wavelength range when stimulated by X-rays, and a photo-detecting component configured to generate a first signal and a second signal, wherein the first signal and second signals are substantially linear functions of the number of photons emitted within the first and second wavelength ranges. Further prior art is known from <CIT>, <CIT> and <CIT>. Positron emission tomography (PET) is a modality of nuclear medicine for imaging metabolic processes using gamma photons emanated from radiopharmaceuticals ingested by a patient or injected into a patient. Multiple PET images are taken in multiple directions to generate/reconstruct a <NUM>-dimensional PET image and/or multiple slices of a PET image. Before image reconstruction, raw PET data are in projection/sinogram space. PET scanning generally provides useful information regarding the functional condition of the body tissues and systems such as the cardiovascular system, respiratory system, and/or other systems. PET scanning is useful for indicating the presence of soft tissue tumors or decreased blood flow to certain organs or areas of the body.

PET scanners with cylindrical geometry can have a reduction in radial spatial resolution, which increases with increasing distance from the center of the field of view (FOV) of the scanner. The loss is due to a parallax effect, which is in turn due to uncertainty in determining a position of a positron-electron annihilation event ("gamma event") with respect to the line of response (LOR) that joins the scintillators involved in the interaction. When a source is relatively far from the central axis, the difference in position between the true line of flight (LOF) of the photon pair and the estimated LOR can be large. Thus, improvement in depth-of-interaction (DOI) measurement is desirable.

There are several methods to attain DOI information. Many approaches are based on a "stacked geometry" of the PET detector block. Two or more layers of scintillator arrays are provided. The gamma rays from a positron-electron annihilation event are absorbed by one of the scintillators, and cause emission of photons. DOI information can be obtained by determining in which layer of scintillator a gamma ray was actually absorbed.

This description of the exemplary embodiments is intended to be read in connection with the accompanying drawings, which are to be considered part of the entire written description. In the description, relative terms such as "lower," "upper," "horizontal," "vertical,", "above," "below," "up," "down," "top" and "bottom" as well as derivative thereof (e.g., "horizontally," "downwardly," "upwardly," etc.) should be construed to refer to the orientation as then described or as shown in the drawing under discussion. These relative terms are for convenience of description and do not require that the apparatus be constructed or operated in a particular orientation. Terms concerning attachments, coupling and the like, such as "connected" and "interconnected," refer to a relationship wherein structures are secured or attached to one another either directly or indirectly through intervening structures, as well as both movable or rigid attachments or relationships, unless expressly described otherwise.

Radiation detectors (e.g., phoswich detector or Positron emission tomography (PET) block detector) can measure depth-of-interaction (DOI) information for each particle or gamma event. This additional DOI information allows a more accurate determination of the respective line-of-response (LOR) sinogram position for each gamma event. In turn, the achievable spatial resolution improves, and finer structures can be resolved, which can have a positive impact on the diagnostic value of the reconstructed images.

According to one aspect of this disclosure, a positron-emission-tomography (PET) scanner includes a PET block detector capable of measuring depth-of-interaction (DOI). The PET block detector comprises two or more scintillators, stacked one-above-the-other and having respectively different peak wavelengths and/or emission spectrum characteristics. The bottom scintillator is positioned directly on a two-dimensional (2D) array of phoswich detectors, such as silicon photomultipliers (SiPMs), without color filters between the SiPMs and the bottom scintillator. The 2D array can have equal numbers of two types of SiPMs, each type having a different spectral response characteristic from the other, so comparison of the total energy collected by all of the SiPMs of each individual type indicates in which of the scintillators the gamma event occurs. Because there are no color filters between the SiPMs and the scintillators, attenuation of the light rays emitted by the scintillators is avoided, wavelength shifts (and corresponding timing shifts) are avoided, and blurring of the sinogram data is reduced.

In some embodiments, a radiation detector comprises a plurality of scintillators. A first one of the scintillators is positioned over and contacting each of the plurality of photon detectors. Each photon detector is capable of detecting impingement of photons thereon. The plurality of photon detectors include first detectors and second detectors. The first detectors are more sensitive to the distribution of light in first emission spectrum than the second emission spectrum. The second detectors are more sensitive to the distribution of light in the second emission spectrum than the distribution of light in the first emission spectrum. The first detectors are more sensitive to the first peak wavelength than the second peak wavelength. The second detectors are more sensitive to the second peak wavelength than are the first detectors. In some embodiments, the second detectors are more sensitive to the second peak wavelength than first peak wavelength.

The two or more layers of scintillators can be made of two different scintillator materials which have two different peak emission wavelengths. In some embodiments, there are no reflectors between the scintillators. In other embodiments the scintillator layers have reflectors only covering part of the length (in the Z direction) of the scintillators, so that the light is shared across the entire detector block (in the X and/or Y directions).

<FIG> shows a schematic diagram of a positron emission tomography (PET) system <NUM>. The system <NUM> comprises: a tomograph <NUM>, an examination table <NUM> for a patient <NUM> who can be moved on the examination table <NUM> through an opening <NUM> of the tomograph, a control device <NUM>, a processor <NUM> and a drive unit <NUM>. The control device <NUM> activates the tomograph <NUM> and receives from the tomograph <NUM> signals which are picked up by the tomograph <NUM>. With the aid of the tomograph <NUM> positron emission sinogram data can be collected. Also disposed in the tomograph <NUM> is a ring of PET detector blocks 9a, 9b (collectively referred to as <NUM>) for acquiring photons which are created by annihilation of electrons and positrons in the PET detector blocks 9a, 9b. Although only <NUM> detector blocks 9a, 9b are shown in <FIG> for ease of viewing, tomograph <NUM> can have many detector blocks <NUM> arranged around the circumference of the tomograph <NUM>. The control device <NUM> is further operable to receive signals from the detector blocks 9a, 9b and is capable of evaluating these signals for creating positron emission tomography images. The control device <NUM> further activates the drive unit <NUM> in order to move the examination table <NUM> in a direction Z together with the patient <NUM> through the opening <NUM> of the tomograph <NUM>. The control device <NUM> and the processor <NUM> can, for example, comprise a computer system with a screen, a keyboard and a data medium <NUM> on which electronically-readable control information is stored, which is embodied so that it carries out the method described below when the data medium <NUM> is used in the processor <NUM> and the control device <NUM>.

<FIG> is a cross sectional diagram of a radiation detector <NUM> sensitive to gamma radiation γ or other radiation (e.g., X-ray, alpha particles, beta particles). PET detector block <NUM>, which is representative of PET detector blocks 9a and 9b (<FIG>). The detector block <NUM> comprises two or more layers (e.g., crystal layers) <NUM> and <NUM>. A first scintillator layer <NUM> may have one or more scintillators <NUM> of a first scintillation material having a first peak wavelength and first emission spectrum. A second scintillator layer <NUM> is positioned on the first scintillator layer <NUM>. The second scintillator layer <NUM> has a plurality of layers (e.g., crystals) <NUM> of a second scintillation material. The second scintillator has a second peak wavelength and second emission spectrum different from the first peak wavelength and emission spectrum. Scintillator layer <NUM> can have many crystals <NUM>, and Scintillator layer <NUM> can have many crystals <NUM>. For example, <FIG> shows an arrangement with a 5x5 array of crystals <NUM> and a 5x5 array of crystals <NUM>. The number of crystals <NUM>, <NUM> in the scintillator layers <NUM>, <NUM> can be varied (e.g., 3x3, 4x4, 4x8, etc.). The first scintillator layer <NUM> and the second scintillator layer <NUM> can have the same number of crystals as each other.

In some embodiments, there are no reflectors between the crystals <NUM>, <NUM> within each scintillator layer <NUM>, <NUM>. In other embodiments the scintillator layers <NUM>, <NUM> can include reflectors (not shown) which cover only part of the length of each scintillator, such that the light is shared across the entire scintillator layer. In other embodiments (not shown), each scintillator layer <NUM>, <NUM> includes a single monolithic crystal (not shown).

The first and second scintillator materials have different emission characteristic spectrums from each other. The first and second scintillator materials have different emission peak wavelengths from each other, as discussed below with respect to <FIG>. In some embodiments, the first and second scintillator materials emit different colors of light from each other.

Referring again to <FIG>, in some embodiments, the first scintillator layer <NUM> comprises Lutetium-yttrium oxyorthosilicate, Lu<NUM>(<NUM>-x)Y2xSiO<NUM>, (LYSO) and emits blue light upon absorbing high energy photons. The second scintillator layer <NUM> can be green-shifted relative to the first scintillator layer <NUM>. For example, in some embodiments, the second scintillator <NUM> comprises Gd<NUM>Al<NUM>Ga<NUM>O<NUM>, (GAGG), BGO, or garnet structures, and emits green light upon absorbing high energy photons. In other embodiments, the scintillator layers <NUM>, <NUM> can comprise other scintillation materials. In some embodiments, the first scintillation material of scintillator layer <NUM> is NaI(TI) and the second scintillation material of scintillator layer <NUM> is BGO. In other embodiments, the first scintillation material of scintillator layer <NUM> is LSO and the second scintillation material of scintillator layer <NUM> is LuYAP. These are only examples, and are not exclusive.

Directly underneath the first scintillator layer <NUM> is an array <NUM> of two groups of photon detectors, such as SiPMs <NUM> and <NUM>, labeled B and G, respectively. The first scintillator layer is positioned over and contacting each of the plurality of photon detectors <NUM>, <NUM>. Each photon detector <NUM>, <NUM> is capable of detecting impingement of a single photon thereon. The plurality of photon detectors include first detectors <NUM> sensitive to photons of the first peak wavelength and first emission spectrum and second detectors <NUM> sensitive to photons of the second peak wavelength and second emission spectrum.

The first detectors (SiPMs) <NUM> and second detectors (SiPMs) <NUM> are arranged in a two-dimensional (2D) array <NUM> for measuring 2D coordinates of each impinging photon. The SiPMs <NUM>, <NUM> can be arranged in a checkerboard pattern with alternating colors in each direction, as shown in <FIG>, for example. The SiPMs <NUM>, <NUM> do not include color filters, and there are no color filters between the SiPMs <NUM>, <NUM> and the first scintillator layer <NUM>. This configuration avoids attenuation (due to light absorption by color filters), and reduces blurring of the light impinging on the detector array <NUM>.

In the example of <FIG>, the scintillator layers <NUM>, <NUM> have more crystals <NUM>, <NUM> per layer than there are SiPMs <NUM>, <NUM> in the array <NUM>. The PET detector block <NUM> has 5x5 arrays of crystals <NUM>, <NUM> in the scintillator layers <NUM>, <NUM>, and a 4x4 array <NUM> of SiPMs <NUM>, <NUM>. This is only one example and is not limiting. For example, the number of crystals <NUM>, <NUM> in each scintillator layer <NUM>, <NUM> can be equal to, or greater than, the number of SiPMs <NUM>, <NUM> in the array <NUM>.

The difference in the spectral sensitivity of the two kinds of SiPMs <NUM>, <NUM> can be achieved by using intrinsic properties of the semiconductor-based photo sensors (e.g. doping profile, pn junction depth, or front-versus-backside illumination geometry). For example, the first detectors and second detectors can have different doping profiles (for example with different depths of the multiplication zone or doping concentrations) or other micro-cell features.

The doping profile determines the shape of the electric field in a semi-conductor device, such as an avalanche photodiode (APD) used in an SiPM. At the pn-junction there is a strong electric field, which provides the multiplication of the charge carriers (avalanche effect). The multiplication or avalanche region is adjacent to a "drift region" (p-doped) having a lower electric field. If a photon is absorbed in the drift region, the electron can drift to the multiplication region, where it is multiplied. If the photon is absorbed in the multiplication (avalanche) region, it can also be multiplied. So, the SiPM creates a measurable signal, when the incident photon is absorbed in the drift or multiplication region.

The absorption length of photons in silicon is wavelength-dependent. If the pn-junction (and therefore the multiplication and drift region) is deeper in the silicon, red light will have a higher probability of being measured because red photons can penetrate silicon more deeply and blue photons have been already absorbed before reaching the pn-junction. If the pn-junction is closer to the surface, the SiPM will be more blue-sensitive, because the blue light is absorbed closer to the surface and a large portion of the red photons will pass through the pn-junction and drift region without being absorbed.

For example, <FIG> is a diagram of a spectral response <NUM> of a first SiPM to blue light, and a spectral response <NUM> of a second SiPM to red light. Although the first SiPM and second SiPM have distinct peaks, there is an overlapping range <NUM> of wavelengths, in which both the first SiPM and the second SiPM output a non-zero value. As a result, an SiPM can be optimized for detecting blue light, and have a smaller, but non-zero, response to red and green light.

A difference in emission spectra, as shown qualitatively in <FIG>, can be achieved by different pn junction depth or different doping profile (i.e., the profile of dopant concentration as a function of depth is different for each type of SiPM).

Similarly, an APD that is more green-sensitive can be achieved with a different pn-junction depth. The SiPM can be optimized for detecting green light, and have a smaller, but non-zero, response to red and blue light. Thus, APDs having different spectral responses can be provided by using different pn-junction depths below the surface of the semiconductor (e.g., silicon) substrate. SiPMs <NUM> and <NUM> having different spectral responses from each other can be produced by including APDs with respectively different doping profile, pn junction depth, or front-versus-backside illumination geometry in the SiPMs <NUM> and <NUM>.

Although exemplary photon detectors <NUM>, <NUM> are described above, any "phoswich" detector capable of detecting low-intensity, low-energy radiation can be used. Here, the term "phoswich" refers to a detector capable of performing this detection function, and is not limited to phosphor sandwich type phoswich devices. The radiation detector <NUM> can include other phoswich detectors, as appropriate for applications including, but not limited to, PET over SPECT to homeland security.

Referring again to <FIG>, the spectral response of the SiPMs <NUM> is sufficiently different from the spectral response of the SiPMs <NUM>, so that the light emitted by the two scintillator materials in scintillator layers <NUM> and <NUM> can be distinguished from one another. In <FIG>, "B" is used as an acronym for mainly blue-sensitive SiPMs and "G" for mainly green-sensitive SiPMs. However, this is only a non-exclusive example, and the method does not require a narrow region of response just in the blue or green regions.

In some embodiments, SiPMs <NUM> have their peak response to light at or near the peak wavelength emitted by the first scintillator layer <NUM>; and SiPMs <NUM> have their peak response to light at or near the peak wavelength emitted by the second scintillator layer <NUM>. However, the method does not require the SiPMs <NUM>, <NUM> to have their peak responses at the same frequencies as the scintillator layers <NUM>, <NUM>. For example, the array <NUM> can detect in which scintillator layer <NUM>, <NUM> the gamma event occurs if: (a) the SiPMs <NUM> output more total current if a gamma event occurs in the first scintillator layer <NUM> than if a gamma event occurs in the second scintillator layer <NUM>; and (b) the SiPMs <NUM> output more total current if a gamma event occurs in the second scintillator layer <NUM> than if a gamma event occurs in the first scintillator layer <NUM>.

The signals of each event are read out by Anger logic which yields an X and Y position to identify the lateral position of the gamma event within the detector block <NUM>. Anger logic is a procedure to obtain the position of incidence of a photon (in the X, Y plane) on the scintillator crystal <NUM> or <NUM>, which involves connecting the SiPM outputs to a resistive network to obtain only four outputs. With these signals, the position of the scintillation centroid can be obtained.

Additionally, the signals of all SiPMs <NUM> belonging to the first group ("B") are summed to calculate a first energy value. Also, the signals of all SiPMs <NUM> belonging to the second group ("G") are summed to calculate a second energy value. Finally, the DOI information can be calculated based on the ratio between the first and second energy values. The threshold ratio between the energy values of the two different scintillator layers <NUM>, <NUM> is dependent on the two types of scintillator materials.

The layout of photon detector <NUM> is only exemplary and is not limiting. A variety of photon detector examples are shown in <FIG>. In each of the examples of <FIG>, the first and second scintillator materials have different emission characteristic spectrums from each other; the first and second scintillator materials have different emission spectra and different peak wavelengths from each other; and the first scintillator layer <NUM> lies directly on (i.e., contacting) the SiPMs with no intervening color filters.

Some embodiments use two different types of SiPMs, where a first type provides better sensitivity than the second type. Some embodiments can take advantage of the sensitivity characteristics by having a majority of the area of the photon detector occupied by the more sensitive type of SiPMs (e.g., the blue-sensitive SiPMs comprising LYSO) and having a minority of the area of the photon detector occupied by the less sensitive type (e.g., green-sensitive) of SiPMs or smaller green-sensitive SiPM chips. This can be accomplished by using fewer second detectors (e.g., <FIG>) or smaller second detectors (e.g., <FIG>) compared to the photon detector <NUM> (<FIG>). This way, the packing fraction of the more sensitive SIPM is increased, and the overall performance (e.g. timing and energy resolution) is improved. <FIG> show three different arrangements in which the first (B) detectors occupy a greater area than the second (G) detectors.

<FIG> shows an example of a photon detector <NUM>, which can be substituted for the photon detector <NUM> in the radiation detector <NUM> of <FIG> in some embodiments. In the photon detector <NUM>, a respective second detector <NUM> (e.g., green) is located in each of a plurality of corners of the 2D array <NUM>, and a remainder of the plurality of photon detectors comprises first detectors <NUM> (e.g., blue). The photon detector <NUM> of <FIG> can provide better time resolution for the first (blue) detectors <NUM>. The photon detector <NUM> of <FIG> can provide good differentiation between the first detectors <NUM> (B) and the second detectors <NUM> (G), due to the high probability of detecting blue light.

<FIG> shows an alternative layout of a photon detector <NUM>, which can be substituted for the photon detector <NUM> in the radiation detector <NUM> of <FIG> in some embodiments. In the photon detector <NUM>, the 2D array <NUM> has a plurality of rows and a plurality of columns. Each of the plurality of rows includes at least one first detector <NUM> (B) and at least one second detector <NUM> (G). Each of the plurality of columns includes at least one first detector <NUM> (B) and at least one second detector <NUM> (G). In some embodiments, as shown in <FIG>, each of the corners has a first detector <NUM> (B), each of the plurality of rows includes a second detector <NUM> (G), and each of the plurality of columns includes a single second detector <NUM> (G). In some embodiments, each row has a different arrangement of first and second detectors from each other row, and each column has a different arrangement of first and second detectors from each other column. The photon detector <NUM> of <FIG> can provide better time resolution for the first (blue) detectors <NUM>, and can provide good differentiation between the first detectors <NUM> (B) and the second detectors <NUM> (G).

<FIG> shows an alternative layout of a photon detector <NUM>, which can be substituted for the photon detector <NUM> in the radiation detector <NUM> of <FIG> in some embodiments. In the photon detector <NUM>, the first detectors (SiPMs) <NUM> and second detectors (SiPMs) <NUM> are arranged in a two-dimensional (2D) array <NUM> for measuring 2D coordinates of each impinging photon. The SiPMs <NUM>, <NUM> can be arranged with alternating colors in each direction. Each of the first detectors <NUM> has a first area, each of the second detectors <NUM> has a second area, and the first area is greater than the second area. For example, the first detectors <NUM> can be identical to the first detectors <NUM> of <FIG>. The second detectors <NUM> can be smaller (and hence less sensitive) than the second detectors <NUM> of <FIG>. That is, the total area of one first detector <NUM> plus one second detector <NUM> can be less than the total area of a first detector <NUM> plus a second detector <NUM> of the photon detector <NUM> of <FIG>. The configuration of photon detector <NUM> provides the same resolution and performance for the first (blue) detectors <NUM> as in the photon detector <NUM> of <FIG>. The configuration of photon detector <NUM> also provides good differentiation between the first detectors <NUM> (B) and the second detectors <NUM> (G), even though the smaller second detectors provide lower resolution than the detectors <NUM>.

<FIG> shows an alternative photon detector configuration <NUM>, which can be substituted for the photon detector <NUM> in the radiation detector <NUM> of <FIG> in some embodiments. The photon detector <NUM> can contain fewer SiPMs at a lower total cost. The first detectors <NUM> and second detectors <NUM> can be the same as the first detectors <NUM> and second detectors <NUM> of <FIG>. The floor plan of the photon detector <NUM> is sparsely populated, including a plurality of vacant regions <NUM> without any SiPMs. The 2D array <NUM> has a plurality of sides. Each of the plurality of sides has a border row or a border column adjacent thereto. Each border row has at least one first detector <NUM> and at least one second detector <NUM>, and each border column has at least one first detector <NUM> and at least one second detector <NUM>. Each border row and each border column has at least one cell without a photon detector. In some embodiments, each of the plurality of rows includes at least one first detector <NUM> (B) and at least one second detector <NUM> (G). Each of the plurality of columns includes at least one first detector <NUM> (B) and at least one second detector <NUM> (G). This configuration provides the same differentiation performance for both B and G type SiPMs, and thus provides the same Z-direction performance for gamma events in the first scintillator layer <NUM> and the second scintillator layer <NUM> for determining the DOI.

<FIG> shows an alternative layout of a photon detector <NUM>, which can be substituted for the photon detector <NUM> in the radiation detector <NUM> of <FIG>. The photon detector <NUM> can have the same arrangement of first detectors <NUM> and second detectors <NUM> as the arrangement of first detectors <NUM> and second detectors <NUM> in photon detector <NUM> of <FIG>. Additionally, all of the vacant regions <NUM> of photon detector <NUM> are replaced with third detectors <NUM> (labeled "S"). Each of the plurality of sides has a border row or a border column adjacent thereto. Each border row has at least one first detector <NUM> and at least one second detector <NUM>, and each border column has at least one first detector <NUM> and at least one second detector <NUM>. Each border row and each border column has at least one cell with a third detector <NUM>. The array has four corners, and each of the four corners has a third detector <NUM>.

Each third detector <NUM> is sensitive to photons in a wavelength band including at least the first peak wavelength (e.g., blue) and the second peak wavelength (e.g., green). In some embodiments, the third detectors <NUM> are SiPMs which are selected to improve timing and energy resolution. For example, the third detectors <NUM> can be SiPMs which have a broad response band, and can capture both blue light and green light, for example. In other examples, the third detectors <NUM> can be "standard" SiPMs (labeled "S") which have a broad response band, and can capture red, green and blue light, for example. The standard SiPMs <NUM> provide high timing and energy resolution in the X-Y plane for each slice, while the first (e.g., B) detectors <NUM> and second (e.g., G) detectors <NUM> differentiate the color of scintillation light for DOI measurement in the Z direction. The broad response band of the standard SiPMs <NUM> allow the SiPMs <NUM> to collect larger amounts of energy, improving resolution and signal-to-noise ratio (SNR) when compared to narrow response band SiPMs.

<FIG> show two alternative photon detector configurations <NUM> and <NUM>, respectively, which can be substituted for the photon detector <NUM> in the radiation detector <NUM> of <FIG> in some embodiments. Out of these two alternatives only the one shown in <FIG> is according to the claimed invention. In some embodiments, the 2D arrays <NUM>, <NUM> have at least one row or column containing at least one split cell <NUM>, <NUM>, and each of the split cells <NUM>, <NUM> comprises one of the first detectors <NUM>, <NUM> having a first area, one of the second detectors <NUM>, <NUM> having a second area. The remaining areas <NUM>, <NUM> are larger than the first area and larger than the second area. In <FIG>, the first detectors <NUM>, <NUM> and second detectors <NUM>, <NUM> are smaller SiPMs, optimized for differentiation of the color of the scintillation light.

<FIG> shows a photon detector <NUM>. The 2D array <NUM> has a plurality of sides. Each of the plurality of sides has a border row or a border column adjacent thereto. Each border row has at least one first detector <NUM> and at least one second detector <NUM>, and each border column has at least one first detector <NUM> and at least one second detector <NUM>. Each border row and each border column has at least one vacant region <NUM> without a photon detector. In some embodiments, each row and each column have at least one split cell <NUM>. Each of the plurality of rows includes at least one first detector <NUM> (B), at least one second detector <NUM> (G), and at least one vacant region <NUM> without a photon detector. The vacant region <NUM> is larger than a sum of the first area of first detector <NUM> and the second area of second detector <NUM>. Each of the plurality of columns has at least one first detector <NUM> (B), at least one second detector <NUM> (G), and at least one vacant region <NUM> without a photon detector. The smaller first detectors <NUM> and second detectors <NUM> can be used as a "binary switch" to generate signals which allow the discrimination between the upper layer <NUM> and lower layer <NUM> of scintillators. The photon detector <NUM> uses sparse SiPM density to reduce cost, while providing binary DOI information.

<FIG> shows a photon detector <NUM> having a layout similar to the layout of <FIG>, except all of the vacant regions <NUM> in <FIG> are replaced with third SiPMs <NUM> in <FIG>. The 2D array has at least one row or column of split cells <NUM>, and each of the split cells <NUM> comprises one of the first detectors <NUM> having a first area, one of the second detectors <NUM> having a second area, and a remaining area that is larger than the first area and larger than the second area. The remaining area includes a third detector <NUM>, which can be an SiPM. The third detector <NUM> is sensitive to photons in a wavelength band including at least the first peak wavelength (of the first scintillators <NUM>) and the second peak wavelength (of the second scintillators <NUM>). In some embodiments, the first peak wavelength corresponds to blue light, the second peak wavelength corresponds to green light, and the third detector is responsive to green light and blue light.

In some embodiments, the third detectors <NUM> are SiPMs which are selected to improve timing and energy resolution, and the first detectors <NUM> and second detectors <NUM> are smaller SiPMs which are optimized for the differentiation of the color of the scintillation light. The smaller "B"- and "G"-SiPMs can be used as a "binary switch" to generate signals which allow the discrimination of whether the scintillation light originates from the lower or upper layers <NUM>, <NUM> of scintillators. The Anger position and timing can either be based on signals from all of the first, second and third detectors <NUM>, <NUM>, <NUM>, or just the third detectors <NUM>. In some embodiments, the remaining area containing the third detector <NUM> is larger than a sum of the first area of the first detectors <NUM> and the second area of the second detectors <NUM>.

The third sensors <NUM> can be SiPMs which have a broad response band, and can capture and detect both blue light and green light, for example. In other examples, the third detectors <NUM> can be capable of detecting red, green and blue light. The standard SiPMs <NUM> provide high timing and energy resolution in the X-Y plane for each slice, while the first (e.g., B) detectors <NUM> and second (e.g., G) detectors <NUM> differentiate the color of scintillation light for DOI measurement in the Z direction using smaller area sensors.

In the examples above, the first and second detectors are blue and green. In other embodiments, the first and second detectors are red and green. In other embodiments, the first and second detectors are red and blue.

<FIG> and <FIG> show two variations in which a third scintillator layer is added. In each of the examples of <FIG> and <FIG>, the first and second scintillator materials have different emission characteristic spectrums from each other; the first and second scintillator materials have different emission peak wavelengths from each other; and the first scintillator layer <NUM> lies directly on (i.e., contacting) the SiPMs with no intervening color filters.

<FIG> shows an embodiment of a PET detection block <NUM> having three scintillator layers <NUM>, <NUM>, <NUM>. The first scintillator layer <NUM> comprises first scintillators <NUM> of a first scintillator material and second scintillators <NUM> of a second scintillator material, as described above with reference to <FIG>. In <FIG>, the third scintillator layer <NUM> of radiation detector <NUM> can comprise third scintillators <NUM> of a third scintillator material which emits red light when a gamma ray is absorbed in the third scintillator layer <NUM>. Directly underneath the first scintillator layer <NUM> is an array <NUM> of three groups of photon detectors, such as SiPMs, labeled R, G, and B, respectively. The first scintillator layer is positioned over and contacting each of the plurality of photon detectors R, G, and B. Each photon detector is capable of detecting photons. The plurality of photon detectors include first detectors sensitive to photons of the first peak wavelength (e.g., blue), second detectors sensitive to photons of the second peak wavelength (e.g., green), and third detectors sensitive to photons of the third peak wavelength (e.g., red).

The first detectors (SiPMs) B, second detectors (SiPMs) G, and third detectors (SiPMs) R are arranged in a two-dimensional (2D) array <NUM> for measuring 2D coordinates of each impinging photon. The SiPMs B, G, R can be arranged round robin in a pattern with alternating colors in each direction, for example. The configuration of R, G and B SiPMs can be varied according to any of the variations described above in <FIG>. For brevity, descriptions of these variations are not repeated. The three scintillator layers <NUM>, <NUM>, <NUM> and three types of SiPMs permit additional resolution in the Z direction.

<FIG> shows an embodiment of another arrangement for a PET detection block <NUM> referred to as a "staggered crystal". The PET detection block <NUM> includes the first and second scintillator layers <NUM>, <NUM> and detector array <NUM> of the PET detection block <NUM> (<FIG>). The staggered crystal configuration of PET detection block <NUM> further includes a third scintillator layer <NUM>. The third scintillator layer <NUM> is offset in the X (or Y) direction from the second scintillator layer <NUM> by a distance <NUM> that is one half the width <NUM> of a crystal <NUM>.

The structure and operation of the first and second scintillator layers <NUM>, <NUM> and the photon detector array <NUM> are the same as described above with respect to the PET detection block <NUM> of <FIG>, a description of which is not repeated for brevity. Additionally the crystals <NUM> of the third scintillator layer <NUM> can comprise the same second scintillator material as the second scintillator layer <NUM>. The light emitted from the third scintillator layer <NUM> can have the same peak wavelength and same emission spectrum as the light emitted from the second scintillator layer <NUM>. Thus, gamma events occurring in the third scintillator layer <NUM> are detected by the second detectors <NUM>, which are the same type of detectors (e.g., green SiPMs) that detect the light from the scintillators <NUM> of the second scintillator layer <NUM>. Because of the offset <NUM> between the crystals <NUM> of the third scintillator layer <NUM> and the crystals <NUM> of the second scintillator layer <NUM>, light emitted by the third scintillator layer <NUM> can be distinguished from light emitted by the second scintillator layer <NUM> based on spatial information. When a crystal <NUM> in the third scintillator layer <NUM> emits photons, the location of peak intensity is below the centroid of the crystal <NUM>, which is offset by a distance <NUM> from the centroid of light emitted by the crystals <NUM> of the second scintillator layer <NUM>. Thus, the processor can determine whether light is received from the second scintillator layer <NUM> or the third scintillator layer <NUM> based on the location of the peak intensity of light received by the second (e.g., green-sensitive) SiPMs.

<FIG> shows an embodiment of the radiation detector <NUM> comprising a first scintillator 92a having a first emission spectrum and a first peak wavelength, and a second scintillator 92b positioned above the first scintillator 92a. The second scintillator 92b has a second emission spectrum and a second peak wavelength different from the first emission spectrum and the first peak wavelength. The first scintillator 92a is positioned over and contacts each of a plurality of photon detectors <NUM>. The plurality of photon detectors <NUM> includes first detectors G sensitive to photons of a first peak emission wavelength of the first emission spectrum and second detectors B sensitive to photons of a second peak emission wavelength of the second emission spectrum. The first scintillator 92a comprises a first layer <NUM> of a first scintillation material, and the second scintillator 92b comprises a second layer <NUM> of the first scintillation material. The first scintillator layer 92a and second scintillator layer 92b can comprise the same scintillation material as each other. The first scintillator 92a further comprises a wavelength shifter <NUM>, between the first layer <NUM> of the first scintillation material and the second layer <NUM> of the first scintillation material. Thus, the addition of the wavelength shifter <NUM> allows both scintillators 92a, 92b to have the same scintillation material as each other, but still provide distinguishably different emission spectra to the detector array <NUM>. The first detectors G are more sensitive to the first emission spectrum (e.g., green) than the second emission spectrum (e.g., blue). The second detectors B are more sensitive to the second emission spectrum (e.g., blue) than the first emission spectrum (e.g., green.

In operation, when radiation (e.g., gamma rays γ) reaches radiation detector <NUM>, most light emitted by the second scintillator 92b and received by the photosensor <NUM> passes through the wavelength shifter <NUM>, where it is wavelength shifted. The light emitted by the scintillator 92a (adjacent the photosensor <NUM>) does not pass through wavelength shifter <NUM> and has its original emission spectrum.

The wavelength shifter <NUM> can be a layer of photo-fluorescent material which absorbs higher frequency photons and emits lower frequency photons. In some embodiments, the wavelength shifter <NUM> comprises a polyvinyltoluene polymer base. For example, in some embodiments, the wavelength shifter can be a layer of EJ-<NUM> (green) or EJ-<NUM> (red) wavelength shifting plastics, sold by ElJen Technology of Sweetwater, TX. For example, the EJ-<NUM> material has an absorption peak wavelength at <NUM> and an emission peak wavelength at <NUM>. The EJ-<NUM> material has an absorption peak wavelength at <NUM> and an emission peak wavelength at <NUM>.

The shift in the wavelength between absorbed light and emitted light is sufficient to be detected by the plurality of detectors. For example, wavelength shifter <NUM> can shift the peak emission wavelength upwards, from blue to green, from green to red, or from red to infrared. The wavelength shifter <NUM> also changes the spectrum of the light emitted by the second scintillator 92b.

<FIG> show another embodiment of a radiation detector <NUM> where the second detectors <NUM> differ from the first detectors <NUM> in front-versus-backside illumination geometry. The radiation detector <NUM> can have the same scintillators 92a, 92b as the radiation detector <NUM> of <FIG>, and a different array <NUM> of photo detectors. The radiation detector <NUM> includes a first scintillator 92a having a first emission spectrum and a second scintillator 92b positioned above the first scintillator 92a. The first scintillator 92a comprises a first layer <NUM> of a first scintillation material, and the second scintillator 92b can comprise a layer <NUM> of the second scintillation material. The second scintillator 92b has a second emission spectrum different from the first emission spectrum. The first scintillator 92a is positioned over and contacts an array <NUM> containing a plurality of photon detectors <NUM>, <NUM>. The array <NUM> of photon detectors includes first detectors <NUM> sensitive to photons of a first peak emission wavelength of the first emission spectrum and second detectors <NUM> sensitive to photons of a second peak emission wavelength of the second emission spectrum. The first scintillator layer 92a and second scintillator layer 92b can comprise different scintillation materials from each other. so as to provide distinguishably different emission spectra to the detector array <NUM>.

<FIG> is an enlarged detail showing a first detector <NUM> and a second detector <NUM>. In some embodiments, the first detector <NUM> and the second detector <NUM> can have the same type of photo sensor integrated circuit <NUM>, but respectively different front-versus-backside illumination geometries. Each photo sensor integrated circuit <NUM> has a front (active) face <NUM> and a backside face <NUM>. The photo sensor integrated circuits <NUM> are capable of front illumination or backside illumination. For example, in the configuration shown, the first detectors <NUM> are front illuminated, and the second detectors <NUM> are backside illuminated. Light detected by the second (backside illuminated) detectors <NUM> passes through an additional semiconductor layer before impinging on the front face <NUM> of the photo sensor integrated circuits <NUM> of the second detectors <NUM>, shifting the wavelength of the detected light.

Although the first detectors <NUM> and second detectors <NUM> of <FIG> include the same photo sensor integrated circuit <NUM> as each other, in other embodiments (not shown) the first detectors and second detectors may have respectively different front-versus-backside illumination geometries as well as respectively different photo sensor integrated circuits.

<FIG> is a flow chart of a method of determining a depth of interaction for a gamma event in a positron emission tomography system having a plurality of scintillator layers, with the bottom one of the scintillator layers in direct contact with an array of SiPMs, and without any color filter between the SiPMs and the bottom scintillator layer.

At step <NUM>, a tracer is administered to a subject (e.g., by injection). The tracer can be a positron-emitting radionuclide, such as fludeoxyglucose (FDG).

At step <NUM>, a pair of gamma rays emitted by decay of the tracer is detected using two or more stacked scintillator layers in each of a pair of PET detector blocks. Each gamma ray is detected by a respective PET detector block approximately <NUM> degrees apart from each other. One of the scintillator layers is above and contacting an array of filterless silicon photomultipliers (SiPMs). The array has a plurality of first SiPMs with a first spectral response, and a plurality of second SiPMs with a second spectral response different from the first spectral response.

Steps <NUM>-<NUM> determine which of the first SiPMs or the second SiPMs outputs more energy in response to the detecting.

At step <NUM>, the control device <NUM> computes a first sum of energy output by the first SiPMs, which indicates the total light energy emitted by the first scintillator layer <NUM> in response to a gamma event.

At step <NUM>, the control device <NUM> computes a second sum of energy output by the second SiPMs, which indicates the total light energy emitted by the second scintillator layer <NUM> in response to a gamma event.

At step <NUM>, the control device <NUM> computes a ratio of the second sum to the first sum.

At step <NUM>, the control device <NUM> compares the ratio (of the second sum to the first sum) to a threshold value, and identifies a depth of interaction of the gamma rays based on whether the ratio exceeds the threshold. The ratio is based on which of the first SiPMs or the second SiPMs outputs more energy per detector. The threshold can be <NUM> in embodiments where the SiPM array <NUM> has equal numbers of first detectors <NUM> and second detectors <NUM> (e.g., <FIG>). Alternatively, the threshold can be normalized in the case (e.g., <FIG>) where the SiPM array has unequal numbers of first detectors <NUM>, <NUM> and second detectors <NUM>, <NUM>. For example, if there are three times as many first detectors <NUM> as second detectors <NUM>, the threshold ratio can be <NUM>/<NUM>. Similarly, the threshold can be normalized to compensate for different detector areas and/or different absorptivities between the first detectors <NUM> and the second detectors <NUM>.

At step <NUM>, the gamma rays from the gamma event are identified as having been absorbed in the first scintillator layer <NUM> in the case where the ratio is less than or equal to the threshold value. Thus, the DOI is identified as being within the first scintillator layer <NUM>.

At step <NUM>, the gamma rays from the gamma event are identified as having been absorbed in the second scintillator layer <NUM> in the case where the ratio exceeds the threshold value. Thus, the DOI is identified as being within the second scintillator layer <NUM>.

The methods and system described herein may be at least partially embodied in the form of computer-implemented processes and apparatus for practicing those processes. The disclosed methods may also be at least partially embodied in the form of tangible, non-transitory machine readable storage media encoded with computer program code. The media may include, for example, RAMs, ROMs, CD-ROMs, DVD-ROMs, BD-ROMs, hard disk drives, flash memories, or any other non-transitory machine-readable storage medium, wherein, when the computer program code is loaded into and executed by a computer, the computer becomes an apparatus for practicing the method. The methods may also be at least partially embodied in the form of a computer into which computer program code is loaded and/or executed, such that, the computer becomes a special purpose computer for practicing the methods. When implemented on a general-purpose processor, the computer program code segments configure the processor to create specific logic circuits. The methods may alternatively be at least partially embodied in a digital signal processor formed of application specific integrated circuits for performing the methods.

Claim 1:
A radiation detector (<NUM>, <NUM>, <NUM>, <NUM>), comprising:
a first scintillator (<NUM>, 92a) having a first emission peak wavelength;
a second scintillator (<NUM>, 92b) positioned on the first scintillator (<NUM>, 92a), the second scintillator (<NUM>, 92b) having a second emission peak wavelength different from the first emission peak wavelength; and
a plurality of photon detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>), the first scintillator (<NUM>, 92a) positioned over and contacting each of the plurality of photon detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>), the plurality of photon detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) including first detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) and second detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>), where the second detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) differ from the first detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) in doping profile, pn junction depth, or front-versus-backside illumination geometry, the first detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) being more sensitive to the first emission peak wavelength than the second emission peak wavelength, the second detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) being more sensitive to the second emission peak wavelength than are the first detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>),
wherein the first detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) and second detectors (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) are arranged in a two-dimensional (2D) array for measuring 2D coordinates of impinging photons,
wherein the 2D array has a plurality of sides, each of the plurality of sides has a border row or a border column adjacent thereto, and each border row and each border column has at least one first detector (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) and at least one second detector (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>),
characterized in that each border row and each border column has at least one third detector (<NUM>, <NUM>), each third detector (<NUM>, <NUM>) is sensitive to photons in a wavelength band including the first emission peak wavelength and the second emission peak wavelength.