Patent Grant
9075150
This invention generally relates to digital radiographic imaging and more particularly relates to a flat panel imaging apparatus having a scintillating phosphor screen with an imaging array of photosensors and thin-film transistor readout devices formed directly on the scintillating phosphor screen.
Generally, medical X-ray detectors employing a scintillating phosphor screen to absorb X-rays and produce light suffer the loss of spatial resolution due to lateral light diffusion in the phosphor screen. To reduce lateral light diffusion and maintain acceptable spatial resolution, the phosphor screens must be made sufficiently thin.
The spatial resolution and X-ray detection ability of an imaging apparatus are often characterized by the modulation transfer function (MTF) and X-ray absorption efficiency, respectively. Thin phosphor screens produce better MTF at the expense of reduced X-ray absorption. Usually, the coating density and the thickness of the phosphor screen are used in the design tradeoff between spatial resolution and X-ray absorption efficiency.
For example, the Lanex Fine and the Lanex Fast Back screens are two typical commercial screens, both manufactured by Eastman Kodak Co. Both are made of Gd2O2S(Tb) phosphor. The Lanex Fast Back screen is relatively thicker and absorbs X-rays more efficiently, but has lower resolution than the Lanex Fine screen. On the other hand, the Lanex Fine screen is thinner than the Lanex Fast Back screen, absorbs X-rays relatively less efficiently, but has higher resolution. The coating density of the Lanex Fine and the Lanex Fast Back screens are 34 mg/cm2 and 133 mg/cm2, respectively. The Lanex Fine and the Lanex Fast Back screens have X-ray absorption efficiencies of 24% and 63% (for 80 kVp, with tungsten target, 2.5-mm Al inherent filtration, and filtered by 0.5-mm Cu+1.0-mm Al) and MTF values of 0.26 and 0.04 at 5 c/mm, respectively.
Recently, digital flat panel X-ray imagers based upon active matrix thin film electronics have become a promising technology for applications such as diagnostic radiology and digital mammography. There are two types of X-ray energy conversion methods used in digital radiography (DR), namely, the direct and indirect method. In the direct method, the X-rays absorbed in a photoconductor are directly transduced into a charge signal, stored on the pixel electrodes on an active matrix array (AMA) and read out using thin film transistors (TFTs) to produce a digital image. Amorphous selenium (a-Se) is usually used as the photoconductor.
In the indirect method, a single phosphor screen is used to absorb X-rays and the resultant light photons are detected by an AMA with a single photodiode (PD) and a TFT switch at each pixel. The photodiode absorbs the light given off by the phosphor in proportion to the X-ray energy absorbed. The stored charge is then read out, like the direct method, using the TFT switch. Common phosphor materials include powder phosphors such as Gd2O2S(Tb) and structured phosphors such as CsI(Tl). Amorphous hydrogenated silicon (a-Si:H) is commonly used to form the photodiode and the TFT switch in the indirect method.
As shown in
The operation of the a-Si based indirect flat panel imager is known by those skilled in the art, and thus only a brief description is given here. Incident X-ray photons are converted to optical photons in the phosphor screen 12, and these optical photons are subsequently converted to electron-hole pairs within the a-Si:H n-i-p photodiodes 70. In general, a reverse bias voltage is applied to the bias lines 85 to create an electric field (and hence a depletion region) across the photodiodes and enhance charge collection efficiency. The pixel charge capacity of the photodiodes is determined by the product of the bias voltage and the photodiode capacitance. The image signal is integrated by the photodiodes while the associated TFTs 71 are held in a non-conducting (“off”) state. This is accomplished by maintaining the gate lines 83 at a negative voltage. The array is read out by sequentially switching rows of TFTs to a conducting state by means of TFT gate control circuitry. When a row of pixels is switched to a conducting (“on”) state by applying a positive voltage to the corresponding gate line 83, charge from those pixels is transferred along the data lines 84 and integrated by the external charge-sensitive amplifiers 86. The row is then switched back to a non-conducting state, and the process is repeated for each row until the entire array has been read out. The signal outputs from the external charge-sensitive amplifiers 86 are transferred to the analog-to-digital converter (ADC) 88 by the parallel-to-serial multiplexer 87, subsequently yielding a digital image. The flat panel imager is capable of both single-shot (radiographic) and continuous (fluoroscopic) image acquisition.
The conventional scintillating phosphor screen imaging panel has three basic components: a substrate of glass or other rigid, transparent material, a TFT layer formed on the substrate, and a phosphor layer containing the scintillator material. There would be advantages in simplifying the design of the imaging panel and reducing size, weight, and cost by eliminating components that are not directly involved in obtaining the image data.
It is an object of the present invention to provide a projection radiographic imaging apparatus that has a simplified and lightweight design and improved detection and display characteristics.
In one aspect, the present invention relates to a projection radiographic imaging apparatus that includes a scintillator and an imaging array. The imaging array includes a plurality of pixels formed directly on a side of the scintillator. Each of the pixels includes at least one photosensor and at least one readout element.
According to another aspect, the present invention includes a method of making a radiographic imaging device. That method includes a step of forming a release layer on a temporary substrate. The method also includes forming an imaging array including a plurality of photosensors and a plurality of thin-film transistor readout elements on the release layer. A scintillator is formed on the imaging array, and the release layer is activated to remove the array from the temporary substrate.
In yet another embodiment, the present invention provides a radiographic imaging panel that includes a first scintillator having a first thickness and a second scintillator having a second thickness. An imaging array is formed on the first scintillator and disposed between the first and second scintillators. The imaging array includes a plurality of photosensors and a plurality of thin-film transistor readout elements.
According to a still further aspect, the present invention provides a radiographic imaging panel including first and second scintillators, having respective first and second thicknesses, and an imaging array disposed directly on one of the first and second scintillators. The imaging array includes a plurality of pixel elements and is positioned between the first and second scintillators. Each pixel element includes a first photosensor optically coupled to the first scintillator, a second photosensor optically coupled to the second scintillator, and a readout element electrically coupled to the first and second photosensors and disposed directly on one of the first and second scintillators.
These and other objects, features, and advantages of the present invention will become apparent to those skilled in the art upon a reading of the following detailed description when taken in conjunction with the drawings wherein there is shown and described an illustrative embodiment of the invention.
While the specification concludes with claims particularly pointing out and distinctly claiming the subject matter of the present invention, it is believed that the invention will be better understood from the following description when taken in conjunction with the accompanying drawings.
The present description is directed in particular to elements forming part of, or cooperating more directly with, apparatuses and methods in accordance with the invention. It is to be understood that elements not specifically shown or described may take various forms well known to those skilled in the art.
In the description that follows, terms and phrases such as “above” or “on top of” are used in a broad sense, to indicate an arrangement of layers relative to each other. Certainly, an X-ray imaging plate may be exposed in any orientation, where stacked layers extend in generally horizontal, vertical, or oblique directions.
The general approach of the present invention eliminates the need for a separate substrate material in forming an imaging panel for radiographic imaging. In the various embodiments of the present invention, the scintillator material of the imaging panel serves as the substrate on which photosensor and TFT readout elements are formed. That is, a separate substrate layer is not needed; the scintillator material serves as the substrate for an imaging panel.
Referring to
Photosensor 112 can be any of a number of types of devices. For example, photosensor 112 can be a segmented or non-segmented metal-insulating semiconductor (MIS). Alternately, photosensor 112 can be a segmented or non-segmented photodiode or a phototransistor. Photosensors are generally well-known in the art, and the invention is not limited to any specific type of photosensor.
Referring to
Although not illustrated, the pixel element 510 shown in
Other arrangements also are contemplated. For example,
Referring to
Specifically,
In
Thus, embodiments of imaging pixel elements according to the invention have been described. In each of these embodiments, an imaging array, including at least one of a photosensor and a readout element is formed directly on a scintillator. According to these embodiments, there is no need for a substrate, as is conventionally used and upon which the imaging array conventionally is formed. These embodiments may also be used in connection with conventional designs using a substrate. For example, when two scintillator layers are used, components of the imaging array may be formed directly on one of the scintillators, as described herein, and other components of the imaging array may be formed on a substrate, on which the other of the scintillator layers also is disposed. In such an embodiment, the two scintillators with accompanying imaging array components may then be “laminated” after construction. In this manner, the benefits of reducing scatter resulting from the substrate are still obtained with one scintillator, whereas the substrate could still provide some structural stability.
Several preferred embodiments of the invention include two scintillator layers of preferably differing thicknesses (although the scintillator layers could be the same thickness). The approach of these embodiments is to utilize multiple scintillator layers in a DR imaging plate to maximize the somewhat conflicting requirements for signal to noise ratio (SNR) and modulation transfer function (MTF). In the embodiments, the first scintillator layer has thickness t1 that is relatively thinner than thickness t2 of the second scintillator layer. With inherently less optical light diffusion, the thinner scintillator layer is optimized for resolution and MTF. Conversely, the thicker scintillator layer is optimized for SNR. For example, the thickness of one phosphor screen may be 97 μm (having a coating weight of 45.3 mg/cm2 of Gd2O2S:Tb), while the thickness of the other phosphor screen may be 186 μm (having a coating weight of 82.7 mg/cm2 of Gd2O2S:Tb). The thinner phosphor screen may have a light control layer of black, absorptive material and the thicker phosphor screen may have a light control coating of black absorptive material. Using the typical X-ray beam for general radiography (the DN5 RQA5 beam), the spatial frequency at which the MTF would equal 50% (f1/2) is 3.8 c/mm and 2.4 c/mm for the thinner phosphor screen and the thicker phosphor screen, respectively. At the same time, the X-ray absorption efficiency of the thicker phosphor screen is 47% as compared to 29% for the thinner phosphor screen. In practical designs, the MTF of the thinner phosphor screen would exceed the MTF of the thicker phosphor screen such that the spatial frequency at which the MTF is 50% (f1/2) for the first phosphor screen is higher than that for the second phosphor screen by at least 0.5 c/mm. In addition, the X-ray absorption efficiency of the second phosphor screen would exceed that of the first phosphor screen by at least 10%. The imaging array is capable of reading the resulting image from each of phosphor screens, so that the combined image can provide higher quality than is available with conventional DR systems with a single phosphor screen.
The material composition of the phosphor screens useful in the embodiments of the invention can include one or more of Gd2O2S:Tb, Gd2O2S:Eu, Gd2O3:Eu, La2O2S:Tb, La2O2S, Y2O2S:Tb, CsI:Tl, CsI:Na, CsBr:Tl, NaLTl, CaWO4, CaWO4:Tb, BaFBr:Eu, BaFCl:Eu, BaSO4:Eu, BaSrSO4, BaPbSO4, BaAl12O19:Mn, BaMgAl10O17:Eu, Zn2SiO4:Mn, (Zn,Cd)S:Ag, LaOBr, LaOBr:Tm, Lu2O2S:Eu, Lu2O2S:Tb, LuTaO4, HfO2:Ti, HfGeO4:Ti, YTaO4, YTaO4:Gd, YTaO4:Nb, Y2O3:Eu, YBO3:Eu, YBO3:Tb, or (Y,Gd)BO3:Eu, or combinations thereof. The phosphor screens, for instance, can be of the same or of different material composition. For example, the phosphor screens may have the same phosphor material but with different particle size distributions, particle-to-binder ratios, packaging densities, or absorbing dye. The median particle size of phosphor material on the second phosphor screen may be in the range from about 1 to about 5 microns, whereas the median particle size of phosphor material on the first phosphor screen may be in the range from about 6 to about 15 microns.
The atomic number of heavy elements may differ in phosphor screens useful in the embodiments of the invention. For example, for higher X-ray energy absorption, the second phosphor screen may have a composition having an element of higher atomic number than that of the first phosphor screen. In one instance, the second phosphor screen may contain Gd2O2S:Tb while the first phosphor screen may contain Y2O2S:Tb. Gadolinium (Gd) has an atomic number of 64, whereas yttrium (Y) has an atomic number of 39.
Moreover, the spatial frequency response of phosphor screens useful in the embodiments of the invention may be different with the use of different phosphor materials with different structures. For example, the second phosphor screen may comprise a columnar structured phosphor such as CsI:Tl, while the first phosphor screen may comprise a powder phosphor such as Gd2O2S:Tb. When evaporated under appropriate conditions, a layer of CsI will condense in the form of needle-like, closely packed crystallites with high packing density. Such a columnar or needle-like phosphor is well known in the art. See, for example, ALN Stevels et al., “Vapor Deposited CsI:Na Layers: Screens for Application in X-Ray Imaging Devices,” Philips Research Reports 29:353-362 (1974); and T. Jing et al, “Enhanced Columnar Structure in CsI Layer by Substrate Patterning”, IEEE Trans. Nucl. Sci. 39: 1195-1198 (1992). In this form, the spatial frequency response (or resolution) is improved over that for a powder phosphor screen of the same thickness, presumably because the columnar crystallites enhance the forward scattering of the light compared to a powder phosphor screen. These columns can be thought to act like fiber optic light guides such that light photons produced by the absorption of an incident X-ray will be guided toward either end of the pillars. Similar to powder screens, a reflective backing is used to maximize the light collection capabilities of the layer by redirecting light photons toward the exit surface. For example, the second phosphor screen may have a CsI:Tl layer with a thickness of 89 microns, while the first phosphor screen may have a Gd2O2S:Tb layer with a thickness of 93 microns. The spatial frequency response of the second phosphor screen may be higher than that of the first phosphor screen. The values of the spatial frequency at which the MTF equals 50% (f1/2) are 4.7 c/mm and 3.3 c/mm for the second and first phosphor screens, respectively.
In these dual screen devices, x-ray radiation generally is incident on the side of the imaging device having the thinner phosphor screen, i.e., the thinner screen is arranged closer to the x-ray source. In this manner, MTF of the thinner screen is optimized. Alternatively, x-ray radiation may be incident on the thicker of phosphor screens, such that the SNR of the thicker screen is optimized.
Also in the dual screen devices, thicknesses t1, t2 of the scintillator layers may be scaled in order to optimize optical absorption lengths. For example, thicknesses exceeding one or more absorption length can be beneficial for reducing optical crosstalk.
Also in the embodiments described above, the light shield may be metal or some other opaque material and may also provide an electrical connection, such as a contact, for example. The light shield additionally/alternatively may be a colorant and/or a semiconductor.
When more than one photosensor is present in the foregoing embodiments, they may be identical, exhibiting the same overall response to incident radiation of different wavelengths. However, in other embodiments, the photosensors may have different sensitivity characteristics, “tuned” to match the emission characteristics of their corresponding scintillator layer. For example, one scintillator layer may emit light with a peak value near 500 nm and the other scintillator layer, could have a different phosphor material, which may emit light with a peak value near 550 nm, for example.
Fabrication
As noted in the foregoing embodiments, the photosensors and/or the readout elements are disposed directly on the scintillator layer according to the invention, without an intervening substrate. In one method of forming this apparatus, these imaging components are fabricated directly on the scintillator layer using known fabrication techniques. Preferred embodiments of this method include using Gd2O2S:Tb as the scintillator layer. Moreover, to ensure proper formation of the imaging components on the scintillator, generally it is desirable to maintain a planarity of the scintillator layer to within 20 nm RMS.
In another preferred embodiment, the imaging components may be formed first on an intermediate or temporary substrate, then laminated onto the scintillator layer. One method using the intermediate substrate includes forming a release layer on the intermediate substrate. Then the array consisting of photosensors and thin-film transistor readout elements are formed on the release layer. Next, radiant energy, heat, pressure, or other energy is applied to the release layer in order to separate it from the intermediate substrate. The release layer or some other substance or process then may be used to adhere or otherwise bond the array of imaging pixels to the scintillator to form an imaging panel. The release layer used in accordance with the invention may be organic or inorganic and may be chemically-, optically-, or thermally-activated. Known release layers include ProLIFT, commercially available from Brewer Science and HD-3007 polyimide adhesive, a laser-activated release commercially available from HD Microsystems. The release layer also may include polybenzoxazole.
In another method of fabricating arrays on a scintillator, a temporary substrate is provided. Like in the embodiment just described, a release layer is formed on a substantially planar surface of the temporary substrate and the imaging array, including photosensitive elements and switching elements, is formed on the release layer. While the array and release layer are still disposed on the temporary substrate, the scintillator is bonded on the imaging array. The release layer is thereafter activated, to separate the imaging array (with scintillator bonded thereto) from the temporary substrate. Accordingly, an imaging array as exemplified in
The release layer preferably is an organic or an inorganic material that can be thermally-, optically-, and/or chemically-activated. For example, in one embodiment the release layer is activated using etching techniques. In such an embodiment, a barrier layer may be used between the release layer and the imaging array and the etch used has a high selectivity of etch rate to the barrier layer, to avoid damaging the imaging array. The etch could be, for example, a xenon difloride gas. In yet another embodiment, the release layer may include a colorant. To activate the release layer, radiant energy is directed to the release layer and the colorant absorbs at least a portion of the radiant energy. The release layers described above also may be used in this embodiment.
While the barrier layer may be include to act as an etch stop, a barrier layer may be formed between the imaging array and the scintillator, to form a device such as that illustrated in
Variations of the foregoing methods also are contemplated and can be used to form the preferred imaging arrays of the present invention. For example, once the imaging array and scintillator are removed from the temporary substrate, a second scintillator can be bonded on the surface of the imaging array exposed upon such removal. Preferably, the second scintillator has properties different from the properties of the first scintillator. Adding the second scintillator in this manner will form devices such as those shown in
The invention has been described in detail with particular reference to certain preferred embodiments thereof, but it will be understood that variations and modifications can be effected within the scope of the invention as described above, and as noted in the appended claims, by a person of ordinary skill in the art without departing from the scope of the invention.
For example, although photodiodes were used as the exemplary light sensing elements in the foregoing embodiments, the invention is not limited in this regard. Any photosensitive element could be used, including, but not limited to, metal-insulating-semiconductors, p-n junction photodiodes, PIN photodiodes, pinned photodiodes, charge-injection-devices, charge-coupled devices, and phototransistors. The invention also is not limited TFTs as the readout elements. Any of MOS transistors, bipolar transistors, diode switches, charge-injection-devices, and charge-coupled devices could be used. Moreover, the imaging array may be formed in any of amorphous silicon, polycrystalline silicon, single-crystal silicon, organic semiconductors, and one or more of binary, ternary, or quaternary semiconductors containing one of indium, zinc, oxygen, and gallium, or any combination thereof.
Thus, what is provided is an imaging array of photosensors and thin-film transistor readout devices formed directly on a scintillating phosphor screen.
| Number | Name | Date | Kind |
|---|---|---|---|
| 7050538 | Tashiro et al. | May 2006 | B2 |
| 7567646 | Buchinsky | Jul 2009 | B2 |
| 7968853 | Altman et al. | Jun 2011 | B2 |
| 20020162965 | Okada et al. | Nov 2002 | A1 |
| 20060151708 | Bani-Hashemi et al. | Jul 2006 | A1 |
| 20070096033 | Freund et al. | May 2007 | A1 |
| 20070120062 | Li et al. | May 2007 | A1 |
| 20080011960 | Yorkston et al. | Jan 2008 | A1 |
| 20080245968 | Tredwell | Oct 2008 | A1 |
| 20090134334 | Nelson | May 2009 | A1 |
| 20100282973 | Schulz | Nov 2010 | A1 |
| Number | Date | Country |
|---|---|---|
| WO 2006114716 | Nov 2006 | WO |
| WO2008051390 | May 2008 | WO |
| WO2009060968 | May 2009 | WO |
| Entry |
|---|
| IEEE Transactions on Nuclear Science, “Enhanced Columnar Structure in CsI Layer by Substrate Patterning”, Jing et al., vol. 39, No. 5, Oct. 1992, pp. 1195-1198. |
| Philips Res. Repts, “Vapour-Deposited CsI: Na Layers, II Screens for Application in X-Ray Imaging Devices”, A.L.N. Stevels et al., vol. 29, pp. 353-362, 1974. |
| Number | Date | Country | |
|---|---|---|---|
| 20110024642 A1 | Feb 2011 | US |
