Notice of Pre-AIA or AIA Status
The present application, filed on or after March 16, 2013, is being examined under the first inventor to file provisions of the AIA .
Claim Rejections - 35 USC § 112
The following is a quotation of 35 U.S.C. 112(b):
(b) CONCLUSION.- -The specification shall conclude with one or more claims particularly pointing out and distinctly claiming the subject matter which the inventor or a joint inventor regards as the invention.
The following is a quotation of 35 U.S.C. 112 (pre-AIA ), second paragraph:
The specification shall conclude with one or more claims particularly pointing out and distinctly claiming the subject matter which the applicant regards as his invention.
Claims 1-15 are rejected under 35 U.S.C. 112(b) or 35 U.S.C. 112 (pre-AIA ), second paragraph, as being indefinite for failing to particularly point out and distinctly claim the subject matter which the inventor or a joint inventor, or for pre-AIA the applicant regards as the invention. There is insufficient antecedent basis for limitation “ first detector data that is sufficient for reconstruction of an evaluable computed tomography”, or regarding claim 5, “spatial frequencies that satisfy a spatial frequency”, or regarding claim 7, “the first data stream only contains partial data for some less than all of the energy”. It is not clear how much data is sufficient or insufficient. Is it based on image intensity, brightness, size, or histogram or resolution? There is insufficient antecedent, how to perform preparation. The scope of claims does not specify details of how sufficient in the claims are enabled or tied to the field of technology, in such a way as to enable one skilled in the art to which it pertains, to make or use the invention.
DETAILED ACTION
Claim Rejections - 35 USC § 103
The following is a quotation of 35 U.S.C. 103(a) which forms the basis for all obviousness rejections set forth in this Office action:
A patent for a claimed invention may not be obtained, notwithstanding that the claimed invention is not identically disclosed as set forth in section 102, if the differences between the claimed invention and the prior art are such that the claimed invention as a whole would have been obvious before the effective filing date of the claimed invention to a person having ordinary skill in the art to which the claimed invention pertains. Patentability shall not be negated by the manner in which the invention was made.
Claims 1-2, 5-11, 13 and 16, are rejected under 35 U.S.C. 103(a) as being unpatentable over Kim et al (U.S. Pub No: 2023/0277152 A1) in view of Okajima et al (U.S. Pub No: 2022/0202383 A1).
Regarding claim 1, Kim discloses a method for operating a computed tomography facility having a rotatable portion with at least one X-ray detector and a portion fixed relative to the rotatable portion with a computing facility configured to process detector data recorded with the at least one X-ray detector (see abstract, also page 5, paragraph, [0065], there is provided a computed tomography imaging system including a “gantry including a rotating” member on a rotating side, a stationary member on a stationary side, and a data communication system. The rotating member on the rotating side includes an X-ray source configured to emit X-rays, an X-ray detector configured to generate detector data, a data storage unit configured to store the detector data, and processing circuitry configured to process at least part of the stored detector data to generate a processed data set. The stationary member on the stationary side is communicatively coupled to the rotating member on the rotating side, and the data communication system is configured to transfer the processed data set from the rotating member on the rotating side to the stationary member on the stationary side. [0065] As the number of electrons and holes from one x-ray event is proportional to the energy of the X-ray photon, the total charge in one induced current pulse is proportional to this energy. The pulse amplitude can then be measured by comparing its value with one or more thresholds in one or more comparators, and counters are introduced by which the number of cases when a pulse is larger than the threshold value may be “recorded”. In this way it is possible to count and/or “record” the number of X-ray photons with an energy exceeding an energy corresponding to respective threshold value (THR) which has been detected within a certain time frame);
wherein the at least one X-ray detector is assigned a preparation facility for ascertaining detector data to be transmitted to the computing facility via a communication path from raw data of the at least one X-ray detector and the communication path has a wireless communication link with a maximum transmission rate, and wherein, during an imaging process, the method comprises (see above, also pages 2-3, paragraphs, [0023] and [0049 the CT imaging system comprises a rotating side comprising an X-ray source configured to emit X-rays, an X-ray detector, a data storage unit, processing circuitry and a stationary side communicatively coupled to the rotating side “via a data communication” system. The method comprises generating detector “data via” the X-ray detector and storing the detector data in the data storage unit. The method further comprises processing the stored detector data, in the processing circuitry, to generate a processed data set. The method further comprises transferring the processed data set from the rotating side to the stationary side via the data communication system. A commonly used X-ray imaging system is a CT imaging system, which may include an X-ray source or X-ray tube that produces a fan beam or cone beam of X-rays and an opposing array of X-ray detectors measuring the fraction of X-rays that are transmitted through a patient or object. The X-ray source or X-ray tube and detector array are mounted in a gantry 15 that rotates around the imaged object (preparation facility).
Also 4, paragraph, [0058] when using several different threshold values, a so-called energy-discriminating photon-counting detector is obtained, in which the detected photons can be sorted into energy bins corresponding to the various threshold values. Sometimes, this type of photon-counting detector is also referred to as a multi-bin detector. In general, the energy information allows for new kinds of images to be created, where new information is available and image artifacts inherent to conventional technology can be removed. In other words, for an energy-discriminating photon-counting detector, the pulse heights are compared to a number of programmable thresholds (Ti-TN) in the comparators and are classified according to pulse-height, which in turn is proportional to energy. In other words, a photon counting detector comprising more than one comparator is here referred to as a multi-bin photon counting detector. In the case of multi-bin photon counting detector, the photon counts are stored in a set of counters, typically one for each energy threshold. For example, counters can be “assigned” to correspond to the highest energy threshold that the photon pulse has exceeded. In another example, counters keep track of the number of times that the photon pulse cross each energy threshold.
Finally, page 5, paragraphs, [0068] and [0070], high spatial resolution; less sensitivity to electronic noise; good energy resolution: and material separation capability (spectral imaging ability). However, energy integrating detectors have the advantage of high count-rate tolerance. The count-rate tolerance comes from the fact/recognition that, since the total energy of the photons is measured, adding one additional photon will always increase the output signal (within reasonable limits), regardless of the amount of photons that are currently being registered by the detector. This advantage is one of the main reasons that energy integrating detectors are the standard for medical CT today. By applying an electric field over the detector material, the charge carriers are collected by electrodes attached to the detector material. The signal is routed from the detector elements to inputs of parallel processing circuits, e.g., ASICs. It should be understood that the term Application Specific Integrated Circuit, ASIC, is to be interpreted broadly as any general circuit used and configured for a specific application. The ASIC processes the electric charge generated from each X-ray and converts it to digital data, which can be used to obtain measurement data such as a photon count and/or estimated energy. In one example, the ASIC can process the electric charge such that a voltage pulse is produced with maximum height proportional to the amount of energy deposited by the photon in the detector material);
dividing, by a selection unit of the preparation facility, the detector data into: (1) a first data stream with first detector data that is sufficient for reconstruction of an evaluable computed tomography image dataset after completion of raw data acquisition with the at least one X-ray detector with a data transmission rate corresponding at most to the maximum transmission rate; and (2) a second data stream comprising remaining detector data as second detector data (see page 2, paragraphs, [0014-0015] CT imaging systems with a rotating section, such as a rotating member of a gantry, typically send all acquired data through data slip rings from the rotating section to a stationary computer (computing facility), wherein the data is later processed in the stationary computer in order to “reconstruct images” of the subject or object. Development in the CT X-ray imaging field makes increasingly high gantry rotation speeds and higher spatial resolution of the detectors possible; with this, the requirements for sufficiently handling the data increases.
Also, page 3, paragraph, [0048] in communication with and electrically coupled to the analog processing circuitry 25 is an imaging processing system 30, which may include digital processing circuitry 40 and/or a computer 50, which may be configured to perform image reconstruction based on the image data from the X-ray detector. The image processing system 30 may, thus, be seen as the computer 50, or alternatively the combined system of the digital processing circuitry 40 and the computer 50, or possibly the digital processing circuitry 40 by itself if the digital processing circuitry is further specialized also for image processing and/or reconstruction.
Also, page 6, paragraphs, [0078-0080] FIG. 6 is a schematic diagram illustrating an example of a semiconductor detector sub-module according to an exemplary embodiment. This is an example of a semiconductor detector sub-module with the semiconductor sensor “split into detector” elements or pixels 22, where each detector element (or pixel) is normally based on a diode having a charge collecting electrode as a key component. The X-rays enter through the edge of the semiconductor sensor. FIG. 7 is a schematic diagram illustrating an example of semiconductor detector sub-module according to another exemplary embodiment. In this example, the semiconductor sensor 21 is also split into a plurality of so-called depth segments or detector elements 22 in the depth direction, again assuming the X-rays enter through the edge. Normally, a detector element is an individual X-ray sensitive sub-element of the detector. In general, the photon interaction takes place in a detector element and the thus generated charge is collected by the corresponding electrode of the detector element. And, the ASICs may be configured for connection to digital processing circuitry and/or memory circuits or components located outside of the MCM and finally the data will be used as input for reconstructing an image.
Finally, page 10, paragraph, [0129] In FIG. 11 the X-ray detector 114 is an energy-discriminating detector comprising individual detector elements, 114a/114b/114c/114d. The CT imaging system may discern the detector data coming from a specific detector element, store the detector data without losing the information on which detector element it originates from. The stored detector data may be processed by the processing circuitry, by performing one or more operations. For example, one operation may be that detector data from different detector elements are averaged, or that the data from certain elements, e.g., neighboring elements, are grouped together/accumulated. The CT imaging system provide the option of individually storing, processing and/or transferring, between the rotating side and the stationary side, detector data, and/or a processed data set from one or more detector elements 114a/114b/114c/114d. The CT imaging system 100 of FIG. 11 may also comprise a display coupled to the second data communication unit 130-2 and is configured to receive and display the processed data set);
transferring the first data stream as a real-time transmission directly to the computing facility via the communication path; storing the second detector data temporarily in a temporary storage facility (see page 2, paragraphs, [0023-0025], an X-ray detector, a data storage unit, processing circuitry and a stationary side communicatively coupled to the rotating side via a data communication system. The method comprises generating detector data via the X-ray detector and storing the detector data in the data storage unit. The method further comprises processing the stored detector data, in the processing circuitry, to generate a processed data set. The method further comprises transferring the processed data set from the rotating side to the stationary side via the data communication system. [0024] Hence, the first and the second aspects of the present invention share a common general inventive concept of providing an improved handling of the data processing and/or data transferring performed by CT imaging systems. Namely, by storing data on a rotating side of a CT imaging system, processing at least a part of the data stored on the rotating side in order to generate a processed data set on the stationary side, and transferring or sending the processed data set to the stationary side, in order to get a result of a scan or at least relevant data in a CT imaging system on the stationary side in a faster, more efficient and more versatile manner. The present invention provides an adaptive procedure, according to which the amount of data to transfer from the rotating side to the stationary side can be selectively adapted.
Also, page 7, paragraph, [0092] by storing data on a rotating side of a CT imaging system, processing at least a part of the data to generate a processed data set on the rotating side and transferring or sending the processed data set to the stationary side, it is possible to get relevant data at the stationary side in a fast, efficient and versatile manner. Processing at least part of the detector data on the rotating side may reduce the data size of the data being transferred via the data communication system, e.g. the slip ring, by factors of 2-20000×, making many calibration and/or imaging steps capable of sending the data within existing slip-ring bandwidth, thus speeding up the overall process of imaging and/or calibrating. The data storage unit may comprise a dedicated large memory, for example, non-volatile memory express (NVMe). The data storage unit on the rotating side may comprise, in addition or alternatively, a temporary memory needed during a continuous data processing, for example, ASIC, field-programmable gate array (FPGA) register or memory. The data storage unit may comprise, in addition or alternatively, a RAM (random access memory). The processing of the data stored on the rotating side may include one or more operations. As an example, an operation may be for example, but not limited to, X-ray source or X-ray tube spit correction, data accumulation, and pileup correction);
and transmitting the second detector data to the computing facility via the communication path after completion of the raw data acquisition with the at least one X-ray detector and the transferring of the first data stream (see above, also page 2, paragraphs, [0023-0024], and paragraph, [0046}, an X-ray detector, a data storage unit, processing circuitry and a stationary side communicatively coupled to the rotating side via a data communication system. The method comprises generating detector data via the X-ray detector and storing the detector data in the data storage unit. The method further comprises processing the stored detector data, in the processing circuitry, to generate a processed data set. The method further comprises transferring the processed data set from the rotating side to the stationary side via the data communication system. Hence, the first and the second aspects of the present invention share a common general inventive concept of providing an improved handling of the data processing and/or data transferring performed by CT imaging systems. Namely, by storing data on a rotating side of a CT imaging system, processing at least a part of the data stored on the rotating side in order to generate a processed data set on the stationary side, and transferring or sending the processed data set to the stationary side, in order to get a result of a scan or at least relevant data in a CT imaging system on the stationary side in a faster, more efficient and more versatile manner. FIG. 2 is a schematic diagram illustrating an example of a CT imaging system 1 comprising an X-ray source 10, which emits X-rays, an X-ray detector system 20 with an X-ray detector, which detects X-rays after they have passed through the object, analog processing circuitry 25, which processes the raw electrical signals from the X-ray detector and digitizes it, digital processing circuitry 40, which may carry out further processing operations on the measured data, such as applying corrections, storing it temporarily, or filtering, and a computer 50, which stores the processed data and may perform further post-processing and/or image reconstruction).
Regarding claim 1, Kim discloses (see page 2, paragraphs, [0023] the CT imaging system comprises a rotating side comprising an X-ray source configured to emit X-rays, an X-ray detector, a data storage unit, processing circuitry and a stationary side communicatively coupled to the rotating side via a data communication system. The method comprises generating detector data via the X-ray detector and storing the detector data in the data storage unit. The method further comprises processing the stored detector data, in the processing circuitry, to generate a processed data set. The method further comprises transferring the processed data set from the rotating side to the stationary side via the data communication system. [0024] Hence, the first and the second aspects of the present invention share a common general inventive concept of providing an improved handling of the data processing and/or data transferring performed by CT imaging systems. Namely, by storing data on a rotating side of a CT imaging system, processing at least a part of the data stored on the rotating side in order to generate a processed data set on the stationary side, and transferring or sending the processed data set to the stationary side, in order to get a result of a scan or at least relevant data in a CT imaging system on the stationary side in a “faster, more efficient” and more versatile manner.
Also, page 5, paragraphs, [0068-0070] there are many benefits of photon-counting detectors including, but not limited to: high spatial resolution; less sensitivity to electronic noise; good energy resolution: and material separation capability (spectral imaging ability). However, energy integrating detectors have the advantage of high count-rate tolerance. The count-rate tolerance comes from the fact/recognition that, since the total energy of the photons is measured, adding one additional photon will always increase the output signal (within reasonable limits), regardless of the amount of photons that are currently being registered by the detector. When a photon interacts in a semiconductor material, a cloud of electron-hole pairs is created. By applying an electric field over the detector material, the charge carriers are collected by electrodes attached to the detector material. The signal is routed from the detector elements to inputs of parallel processing circuits, e.g., ASICs. The ASIC processes the electric charge generated from each X-ray and converts it to digital data, which can be used to obtain measurement data such as a photon count and/or estimated energy. In one example, the ASIC can process the electric charge such that a voltage pulse is produced with maximum height proportional to the amount of energy deposited by the photon in the detector material.
Finally, page 7, paragraph, [0092] by storing data on a rotating side of a CT imaging system, processing at least a part of the data to generate a processed data set on the rotating side and “transferring or sending” the processed data set to the stationary side, it is possible to get relevant data at the stationary side in a fast, efficient and versatile manner. Processing at least part of the detector data on the rotating side may reduce the data size of the data being transferred via the data communication system, e.g. the slip ring, by factors of 2-20000×, making many calibration and/or imaging steps capable of sending the data within existing slip-ring bandwidth, thus “speeding up the overall process” of imaging and/or calibrating. The data storage unit may comprise a dedicated large memory, for example, non-volatile memory express (NVMe). The data storage unit on the rotating side may comprise, in addition or alternatively, a temporary memory needed during a continuous data processing, for example, ASIC, field-programmable gate array (FPGA) register or memory. The data storage unit may comprise, in addition or alternatively, a RAM (random access memory).
But does not explicitly state, “data transmission rate corresponding to the maximum transmission rate”.
However, Okajima, in the same field of “X-ray detector and computed tomography apparatus and a control method”, teaches (see page 3, paragraphs, [0038-0040] the memory is implemented by a semiconductor memory device such as a random-access memory (RAM) or a flash memory, a hard disk, an optical disk, etc. For example, the memory stores projection data and reconstructed image data. For example, the memory stores detection data under each scan condition and estimated data (estimated count number) of count results under each scan condition obtained during a preliminary collection process. For example, the memory stores various programs. A storage area of the memory may be in the X-ray CT apparatus or in an external storage device connected via a network. Here, the memory 41 is an example of a storage unit. The display presents various types of information. For example, the display outputs a medical image (CT image) generated by the processing circuitry, a graphical user interface (GUI) for receiving various operations from the operator, etc. Any various displays may be used as the display as appropriate. For example, a liquid crystal display (LCD), a cathode ray tube (CRT) display, an organic electro luminescence display (OELD), or a plasma display may be used as the display. The display may be provided at any place in the control room. The display may be provided in the gantry. The display may be of a desktop type or may include a tablet terminal, or the like, which may “wirelessly communicate” with a main body of the console. One or more projectors may be used as the display.
Also, page 6, paragraph, [0075] Here, the transmission data amount per unit time and per detection element increases as the view rate increases. The amounts of data transmitted per detection element in the five energy bins are, for example, as follows.
Also, page 7, paragraphs, [0084-0095] When the estimated count number is used, the bit number per detection element needed for transmission at each view rate (transmission rate), may be estimated as follows. 1000 view/sec: 7.0×10.sup.5 bit/pixel/sec (Reduction rate of the bit number by removing higher-order bits: 18%) 2000 view/sec: 1.3×10.sup.5 bit/pixel/sec (Reduction rate of the bit number by removing higher-order bits: 24%). 3000 view/sec: 1.8×10 bit/pixel/sec. Reduction rate of the bit number by removing higher-order bits: 30%), 4000 view/sec: 2.4×10 bit/pixel/sec (Reduction rate of the bit number by removing higher-order bits: 30%) 5000 view/sec: 2.8×10.sup.5 bit/pixel/sec (Reduction rate of the bit number by removing higher-order bits: 35%). As described above, in a case where the higher-order bits are removed in accordance with the estimated count number, even when the view rate is smaller than 5000 view/sec, the data size transmitted from the X-ray detector 12 to the console 40 may be reduced. According to the above estimation, the transmitted data may be reduced by approximately 20% to 40% in terms of the bit number per detection element needed for transmission at each view rate).
Therefore, it would have been obvious to one having ordinary skill in the art at the time the invention was made to modify Kim invention according to the teaching of Okajima because to combine, the processing system and method of X-ray and Computed Tomography that obtains data transfer rates for raw data to increase specifically “transfer rate” for a transmission and measurement of the generated data as taught by Kim with teaching of Okajima that is in the same field of processing and generating CT data and using wireless connection for measuring the generated data and identify the transmission rate for a more efficient transmission of the data.
Regarding claim 2, Kim discloses the method of claim 1, wherein the first data stream comprises the first detector data that is at least partially reduced in spatial resolution and/or spatial coverage compared to all of the detector data (see claim 1, also page 5, paragraph, [0068] there are many benefits of photon-counting detectors including, high spatial resolution; less sensitivity to electronic noise; good energy resolution: and material separation capability (spectral imaging ability). However, energy integrating detectors have the advantage of high count-rate tolerance. The count-rate tolerance comes from the fact/recognition that, since the total energy of the photons is measured, adding one additional photon will always increase the output signal (within reasonable limits), regardless of the amount of photons that are currently being registered by the detector. This advantage is one of the main reasons that energy integrating detectors are the standard for medical CT today).
Regarding claim 5, Kim discloses the method of claim 1, wherein the preparation facility comprises a transformation unit configured to ascertain a spatial frequency representation of the detector data, and wherein only detector data with spatial frequencies that satisfy a spatial frequency criterion is sorted into the first data stream (see claim 1, also page 9, paragraphs, [0121-0123] FIG. 9 schematically shows a CT imaging system 100 according to an exemplifying embodiment of the present invention. In FIG. 9 the CT imaging system 100 comprises a rotating side 110 and a stationary side 120. FIGS. 3-7, 8a, 8b. The X-ray detector may alternatively be an energy-integrating detector (EID) in which the detected signal is proportional to the total energy deposited by all photons without specific information about an individual photon or its energy. The CT imaging system 100 further comprises a data storage unit 116, which allows detector data from the X-ray detector 114 to be stored on the rotating side 110. The data storage unit 116 may be a non-volatile memory. The CT imaging system 100 further comprises processing circuitry 118 which may receive data, such as detector data, and process it. By the term “processing circuitry” it is here meant circuitry, e.g., a processor, that is capable of performing operations, such as mathematical operations, for processing/transforming data. The processing circuitry may be digital processing circuitry. The processing circuitry 118 and data storage unit 116 may be communicatively connected, such that data may be transferred between them. It is to be understood that the processing circuitry 118 may process detector data and/or already processed data, in order to get a resulting processed data set which may be transferred to the stationary side 120. The processed data set may be a reduced data set, wherein a reduced data set may have a reduced data size. By way of example, the data communication system 130 may comprise a first data communication unit 130-1 on the rotating side and a second data communication unit 130-2 on the stationary side. For example, the data communication system 130 may be a slip ring typically used in CT imaging systems with a rotating member of a gantry. By “slip ring” it is here meant, but not limited to, an electromechanical device that allows the transmission of power and “electrical signals”, e.g., power and data transfer between a rotating structure and a stationary structure. The CT imaging system 100 of FIG. 9 may also comprise a display 132 coupled to the second data communication unit 130-2 and is configured to receive and display the processed data set.
Regarding claim 6, Kim discloses the method of claim 1, wherein, when detector data is omitted spatially, temporally, frequency-based, or a combination thereof for the first data stream, on the reconstruction of a computed tomography image dataset from the detector data of the first data stream alone, at least a part of the omitted detector data is reconstructed by interpolation and/or extrapolation (see claim 1, also pages 8, 11 and paragraphs [0109] and [0149], a first processed data set is transferred via the data communication system, while other parts of the first processed data set is being generated, such that the processing of the detector data and the transferring of a processed data set from the rotating member on the rotating side to the stationary member on the stationary side is performed at least partially simultaneously. The present embodiment allows a CT imaging system to be used in a quicker and more efficient manner, essentially reducing the time it takes from performing the scan of the subject/object to producing an “image reconstructed” from the detector data. FIG. 15, the CT imaging system comprises an X-ray detector, a data storage unit and processing circuitry arranged on the rotating side of the CT imaging system, e.g. on a rotating member of a gantry. Here, the generated detector data from the X-ray detector is stored on the data storage unit and/or processed by the processing circuitry 118 without being pre-processed. The processing circuitry 118 may comprise a logic box 118-1 and/or a processing box 118-2, wherein the processing box 118-2 may be configured to performer higher complexity computations compared to the logic box 118-1. The processing circuitry 118 may process at least part of the detector data on the data storage unit 116, and a data collector 119 may collect data, processed or non-processed, from all channels or a subset thereof and send it to the data communication system 130 which may then transfer the data from the rotating member of the gantry, to a stationary member of the gantry, arranged on the stationary side of the CT imaging system 100. It is further noted that the inventive concepts relate to all possible combinations of features unless explicitly stated otherwise).
Regarding claim 7, Kim discloses the method of claim 1, wherein the computed tomography facility is configured for spectral imaging, wherein the detector data comprises partial data assigned to different energy parameter values, and wherein the first data stream only contains partial data for some less than all of the energy parameter values (see claim 1, also page 4, paragraph, [0058] when using several different threshold values, a so-called energy-discriminating photon-counting detector is obtained, in which the detected photons can be sorted into energy bins corresponding to the various threshold values. Sometimes, this type of photon-counting detector is also referred to as a multi-bin detector. In general, the energy information allows for new kinds of images to be created, where new information is available and image artifacts inherent to conventional technology can be removed. In other words, for an energy-discriminating photon-counting detector, the pulse heights are compared to a number of programmable thresholds (Ti-TN) in the comparators and are classified according to pulse-height, which in turn is proportional to energy. In other words, a photon counting detector comprising more than one comparator is here referred to as a multi-bin photon counting detector. In the case of multi-bin photon counting detector, the photon counts are stored in a set of counters, typically one for each energy threshold. For example, counters can be assigned to correspond to the highest energy threshold that the photon pulse has exceeded. In another example, counters keep track of the number of times that the photon pulse cross each energy threshold).
Regarding claim 8, Kim discloses the method of claim 1, wherein the division is carried out in order to make a greatest possible use of the maximum transmission rate (see claim 1, also page 5, paragraphs, [0067-0070] in a photon counting detector, there is typically a Photon Counting Logic which determines if a new photon has been registered and, registers the photons in counter(s). In the case of a multi-bin photon counting detector, there are typically several counters, for example one for each comparator, and the photon counts are registered in the counters in accordance with an estimate of the photon energy. The logic can be implemented in several different ways. Two of the most common categories of Photon Counting Logic are the so-called non-paralyzable counting modes, and the paralyzable counting modes. Other photon-counting logics include, for example, local maxima detection, which counts, and possibly also registers the pulse height of, detected local maxima in the voltage pulse. There are many benefits of photon-counting detectors including, but not limited to: high spatial resolution; less sensitivity to electronic noise; good energy resolution: and material separation capability (spectral imaging ability). However, energy integrating detectors have the advantage of high count-rate tolerance. The count-rate tolerance comes from the fact/recognition that, since the total energy of the photons is measured, adding one additional photon will always increase the output signal (within reasonable limits), regardless of the amount of photons that are currently being registered by the detector. When a photon interacts in a semiconductor material, a cloud of electron-hole pairs is created. By applying an electric field over the detector material, the charge carriers are collected by electrodes attached to the detector material. The signal is routed from the detector elements to inputs of parallel processing circuits, e.g., ASICs. It should be understood that the term Application Specific Integrated Circuit, ASIC, is to be interpreted broadly as any general circuit used and configured for a specific application. The ASIC processes the electric charge generated from each X-ray and converts it to digital data, which can be used to obtain measurement data such as a photon count and/or estimated energy. In one example, the ASIC can process the electric charge such that a voltage pulse is produced with maximum height proportional to the amount of energy deposited by the photon in the detector material).
Regarding claim 13, Kim discloses the method of claim 1, wherein, during the transmission of the second detector data, initialization information for the next imaging process is already transferred to the rotating portion (see claim 1, also page 8, paragraph, [0108] According to an embodiment, the processing circuitry is configured to process the stored data using a set of consecutive output modes to generate respective processed data sets, and wherein the data communication system is configured to transfer the respective processed data sets consecutively. The present embodiment is advantageous in that the CT imaging system may operate in steps, wherein the steps may be predetermined. For example, this allows the CT imaging system to optimize how the detector data is stored, processed and transferred for first generating an initial image for quick diagnosis, followed by generating a more detailed image for final diagnosis. Furthermore, the present embodiment allows the CT imaging system to process and transfer one or more processed data sets without requiring input from a user, thus reducing downtime for the whole process by avoiding pauses between steps.
Also, page 11, paragraphs, [0136-0137] The CT imaging system 100 may be configured, in e.g., the first output mode for patient imaging, to first produce a less detailed image based on the detector data from a scan of a subject/object, and afterwards produce a more detailed image of the scan. This may be done by initially generating the first processed data set from the first part of the detector data, followed by transferring the first processed data set via the data communication system 130 to the stationary side 120. Subsequently, the CT imaging system 100 may be configured, in the first output mode, to transfer the remaining detector data to the stationary side via the data communication system, without processing it with the processing circuitry. The first output mode of the CT imaging system thus allows an initial, quicker, scan to be performed and shown to a user of the system, while the remaining detector data may be processed and/or transferred afterwards. The CT imaging system may be configured, in e.g., the second output mode, to initially process the second part of the detector data to generate the second processed data set, transfer the second processed data set via the data communication system 130, to the stationary side 120, to provide an image based on the second processed data set. Subsequently, the CT imaging system 100 may be configured to process a third part of the detector data to generate a third processed data set, which is then transferred via the data communication system 130, to the stationary side 120, in order to provide a secondary image with information not present in the second processed data set, e.g., to achieve a more detailed image. The CT imaging system 100 of FIG. 12 may also comprise a display 132 coupled to the second data communication unit 130-2 and is configured to receive and display the processed data set).
With regard to claims 9, 10, 11, 16, the arguments analogous to those presented above for claims 1, 2, 5, 6, 7, 8 and 13 are respectively applicable to claims 9 and 16.
Allowable Subject Matter
Claims 3, 4, 12, 14 and 15 are objected to as being dependent upon a rejected base claim, but would be allowable if rewritten in independent form including all of the limitations of the base claim and any intervening claims.
Contact Information
Any inquiry concerning this communication or earlier communications from the examiner should be directed to Seyed Azarian whose telephone number is (571) 272-7443. The examiner can normally be reached on Monday through Thursday from 6:00 a.m. to 7:30 p.m.
If attempts to reach the examiner by telephone are unsuccessful, the examiner's supervisor, Moyer Andrew, can be reached at (571) 272-9523. The fax phone number for the organization where this application or proceeding is assigned is 571-273-8300.
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/SEYED H AZARIAN/Primary Examiner, Art Unit 2667
January 12, 2026