Method and system for integrated patient table digital X-ray dosimeter

A method of imaging a patient and an X-ray dosimetry system are provided. The X-ray dosimetry system includes a support platform configured to support an object to be imaged and a digital X-ray dosimeter mounted on a surface of the support platform, the X-ray dosimeter configured to receive incident radiation prior to the incident radiation having passed through the object to be imaged, the X-ray dosimeter comprising a thickness of less than about four millimeters.

BACKGROUND

This description relates to radiation detectors, and, more particularly, to a system and method for measuring peak skin dose of a patient directly on a patient examining table.

During fluoroscopic medical imaging procedures, injury to a patient and/or doctor may occur due to excessive exposure to radiation, and skin damage is a risk to both doctors as well as patients. To minimize a dose of radiation during such procedures, it is important to measure the peak dose delivered to the skin during fluoroscopic procedures. However, measurement of the peak skin dose has been challenging, and at least some known imaging systems do not include methods for tracking peak skin dose. Systems which include skin dose monitoring commonly employ calculation and estimation techniques in order to estimate peak skin dose. The estimate involves tracking the patient position and the X-ray tube output to track the peak skin dose delivered to the skin. Those systems do not typically include changes due to an examining table on which a patient is positioned or the patient geometry, for example, a size and a weight of the patient. The examining table may include operating, surgical, or other patient or workpiece support table.

Additionally, dose information for fluoroscopic procedures is also projected or modeled based on the position of the patient and the characteristics of the X-ray tube.

Commercial systems do not exist for the active measurement of peak skin dose across the full area of exposure. Dosimeters are available for personal exposure at a fixed point. Currently dosimeters are available for personal exposure at fixed points and a peak skin dose is estimated from the fixed point exposure, the known properties of the X-ray tube output, the known properties of the examining table on which the patient is positioned, and patient geometry. However, such dose calculations are merely estimates, which are not totalized in real-time.

BRIEF DESCRIPTION

In one embodiment, an X-ray dosimetry system includes a support platform configured to support an object to be imaged and a digital X-ray dosimeter mounted on a surface of the support platform. The X-ray dosimeter is configured to receive incident radiation prior to the incident radiation having passed through the object to be imaged and has a thickness of less than about four millimeters.

In another embodiment, a method of imaging a patient includes providing a patient support table including a digital X-ray dosimeter coupled to a surface of the patient support table. The digital X-ray dosimeter includes a flexible substrate layer, an electrode layer coupled to a conductive interconnect, a photodiode layer, a second electrode layer, and a flexible scintillator. The method further includes positioning a patient on the patient support table with a portion of the patient to be imaged located adjacent to the digital X-ray dosimeter, positioning an imaging detector on a side of the patient opposite from the patient support table, and displaying a peak skin dose to the patient with respect to a position of a body of the patient.

In yet another embodiment, a patient imaging system includes a support platform configured to support a human patient to be imaged and a digital radiation dosimeter configured to receive incident radiation prior to passing through the patient. The digital radiation dosimeter includes a substrate, an electrode electrically coupled to a conductive interconnect, a photodiode layer, and a scintillator layer and a thickness of less than about four millimeters. The system also includes an imaging radiation source positioned on a same side of the patient as the digital radiation detector and an imaging radiation detector positioned on a side opposite of the patient from the imaging radiation source and the digital radiation dosimeter. The digital radiation dosimeter is configured to measure a peak skin dose to the patient accumulated as a function of position of the patient and the imaging radiation source.

DETAILED DESCRIPTION

The following detailed description illustrates embodiments of the disclosure by way of example and not by way of limitation. It is contemplated that the disclosure has general application to structural and methodical embodiments for the active measurement of peak skin dose of radiation to a patient during medical procedures. The measurement system is low cost, thin and flexible, and may be incorporated into an existing framework of surgical systems.

Embodiments of the disclosure describe incorporating a dose measurement device or dosimeter within the examining table that yields a peak skin dose accumulated as a function of position, with a size resolution of approximately 2-3 millimeter (mm) to approximately 1 centimeter (cm). The dosimeter is substantially transparent to the x-ray radiation, stopping only a few percent of the delivered dose. The dosimeter is placed in close proximity to the skin of a patient to accurately measure effects, such as, scattering which are influenced by patient placement, size and examining table geometry. As used herein, peak skin dose refers to a highest dose at any portion of a patient's skin during a procedure (integrated in time at a specific point on the skin).

The peak skin dose measurement device or dosimeter is part of an integrated dose detection system that works by placing a low cost, large area array of high x-ray transparency detectors onto a thin, flexible substrate. The x-ray detector pixels may be formed to be 2-3 mm in size and may be 1 cm or larger. The anode is formed by a conductive metal or conductive oxide. The photosensitive layer is an un-patterned thin film photodiode, which, in various embodiments, is formed of plasma-enhanced chemical vapor deposition (PECVD) amorphous silicon (a-Si) or a solution coated organic photodiode (OPD). A transparent conductive metal or oxide is deposited on top of the OPD absorber layer. The photodiode may alternately be sealed or encapsulated with a transparent material such as glass to form a protective environmental seal for the photodiode. A scintillator material is then attached or deposited onto the top electrode or barrier material. The scintillator thickness is just large enough to generate signal for the minimum detectable dose, which limits the absorption of x-rays in the detector. The scintillator may be a thin film deposited material or a plastic sheet with scintillator particles embedded in plastic binder, laminated using adhesive. Additionally, the scintillator may be directly printed onto the substrate or the photodiode.

The photodiode generates electrical current which increases as a function of x-ray flux during x-ray exposure. The electrical current is conducted thru metallic interconnects which are printed or patterned onto the flexible substrate. The conductors are either on the front (photodiode) side of the substrate and connected to the anode, or on the back side (opposite the photodiode) and electrically connected thru a via in the substrate. At the edge of the substrate, a series of TAB bond pads are provided for connection to external electronics, which are used to amplify and convert the current to digital information. The x-ray dosimeter sheet may be incorporated into existing examining tables either internally or laminated to the surface of the table. In some embodiments, the X-ray dosimeter is formed of a plurality of X-ray dosimeter subassemblies communicatively coupled to a dose measuring circuit that receives and processes signals representative of an amount of radiation received by the X-ray dosimeter subassemblies. In some embodiments, the X-ray dosimeter subassemblies are positioned in an abutting relationship adjacent to each other on the examining table and in other embodiments, the X-ray dosimeter subassemblies are positioned such that a gap is formed between adjacent subassemblies. Moreover, X-ray dosimeter subassemblies are permanently or semi-permanently affixed to the examining table, such as by being formed with the examining table or by using an adhesive. In other embodiments, X-ray dosimeter subassemblies are held in position by gravity or by static attraction.

As used herein, a radiation dosimeter differs from a radiation imaging detector in that a radiation dosimeter is used is an imaging system with an radiation imaging detector to determine peak skin dose with respect to a patient position and a position of a radiation imaging source. The radiation dosimeter is positioned on the same side of the patient as the radiation imaging source and receives the radiation prior to the radiation passing through a body of the patient. The radiation imaging detector receives the radiation after it has passed through the body of the patient. The radiation dosimeter is formed to be as radiation transparent as possible, while the radiation imaging detector is formed to absorb as much of the radiation reaching it as possible. Moreover, the pixel area of the radiation dosimeter is relatively larger than the pixel area of the radiation imaging detector making the resolution of the radiation imaging detector greater than the resolution of the radiation dosimeter.

The following description refers to the accompanying drawings, in which, in the absence of a contrary representation, the same numbers in different drawings represent similar elements.

FIG. 1is a schematic block diagram of an exemplary X-ray imaging system10. In the exemplary embodiment, X-ray imaging system10is configured to acquire and process X-ray image data. X-ray imaging system10includes an X-ray source12, a collimator14, and a detector22. Detector22is a fluoroscopic detector. Alternatively, detector22is a detector that enables the operation of X-ray imaging system10as described herein. In one embodiment, detector22is mounted on a support platform23by either coupling detector22to a surface of support platform23or embedded in a well formed in the surface of support platform23. X-ray source12is positioned adjacent to collimator14. In one embodiment, X-ray source12is a low-energy source and is employed in low energy imaging techniques, such as, but not limited to, fluoroscopic techniques. Collimator14facilitates a stream of X-ray radiation16emitted by X-ray source12to radiate towards a target18, such as an industrial component or a human patient. A portion of X-ray radiation16is attenuated by target18and at least some attenuated radiation20impacts detector22.

Detector22is based on scintillation, i.e., optical conversion, direct conversion, or on other techniques used in the generation of electrical signals based on incident radiation. For example, a scintillator-based detector converts X-ray photons incident on its surface to optical photons. These optical photons may then be converted to electrical signals by employing photosensor(s), e.g., photodiode(s). Conversely, a direct conversion detector directly generates electrical charges in response to incident X-ray photons. The electrical charges are stored and read out from storage capacitors. As described in detail below, these electrical signals, regardless of the conversion technique employed, are acquired and processed to construct an image of the features (e.g., anatomy) within target18.

In the exemplary embodiment, X-ray source12is controlled by a power supply and control circuit24which supplies power and control signals for examination sequences. Moreover, detector22is coupled to a detector acquisition circuit26, which is configured to receive electrical readout signals generated in detector22. Detector acquisition circuit26may also execute various signal processing and filtration functions, such as, for initial adjustment of dynamic ranges and interleaving of digital signals.

In the exemplary embodiment, one or both of power supply/control circuit24and detector acquisition circuit26is responsive to signals from a system controller28. System controller28includes signal processing circuitry, typically based upon a general purpose or application specific digital computer programmed to process signals according to one or more parameters. System controller28also includes memory circuitry for storing programs and routines executed by the computer, as well as configuration parameters and image data and interface circuits.

System10includes a dose measuring circuit30configured to receive acquired radiation data from detector acquisition circuit26. Dose measuring circuit30is configured to process the acquired radiation data to a dose received at detector22. In various embodiments, detector22receives stream of X-ray radiation16before it passes through target18. In these embodiments, X-ray source12and collimator14are located on a side opposite of detector22from target18. For example, inFIG. 1, X-ray source12and collimator14could be positioned below support platform23and configured to direct stream of X-ray radiation16upward through support platform23, detector22, and into target18.

An operator workstation32is communicatively coupled to system controller28and/or dose measuring circuit30to allow an operator to initiate and configure X-ray imaging of target18and to view images generated from X-rays that impinge detector22. For example, system controller28is in communication with operator workstation32so that an operator, via one or more input devices associated with operator workstation32, may provide instructions or commands to system controller28.

Similarly, dose measuring circuit30is in communication with operator workstation32such that operator workstation32receives and displays the output of dose measuring circuit30on an output device34, such as a display or printer. Output device34may include standard or special purpose computer monitors and associated processing circuitry. In general, displays, printers, operator workstations, and similar devices supplied within system10may be local to the data acquisition components or may be remote from these components, such as elsewhere within an institution or hospital or in an entirely different location. Output devices and operator workstations that are remote from the data acquisition components may be operatively coupled to the image acquisition system via one or more configurable networks, such as the Internet or virtual private networks. Though system controller28, dose measuring circuit30, and operator workstation32are shown distinct from one another inFIG. 1, these components may actually be embodied in a single processor-based computing system. Alternatively, some or all of these components may be present in distinct processor-based computing systems configured to communicate with one another. For example, dose measuring circuit30may be a component of a distinct reconstruction and viewing workstation.

FIG. 2is a perspective cut-away view of a physical arrangement of the components of an exemplary scintillation-based detector35suitable for use as detector22depicted inFIG. 1. Detector35includes a flexible substrate36upon which one or more components are deposited. For example, in the present embodiment, detector35includes a continuous photosensor element38, transistors42, such as, but not limited to amorphous Silicon (a-Si), thin-film transistors (TFTs), scintillator44, data readout lines48, scan lines50, a conductive layer54, and a dielectric layer56deposited with respect to substrate36. The components of detector35are composed of metallic, dielectric, organic, and/or inorganic materials, and are fabricated with respect to substrate36using various material deposition and removal techniques. Some examples of deposition techniques include, for example, chemical vapor deposition, physical vapor deposition, electrochemical deposition, stamping, printing, sputtering, and/or any other suitable deposition technique. Some examples of material removal techniques include lithography, etching, such as, but not limited to dry etching, wet etching, laser etching, sputtering, and/or any other suitable material removal techniques.

Detector35includes an array of pixel areas40on flexible substrate36. Each of pixel areas40includes transistors42operatively coupled to respective data readout lines48, scan lines50, and photosensor38. In the present embodiment, transistors42are arranged in a two dimensional array having rows extending along an x-axis51and columns extending along a y-axis52, or vice versa. In some embodiments, transistors42are arranged in other configurations. For example, in some embodiments, transistors42are arranged in a honeycomb pattern. A spatial density of transistors42determines a quantity of pixel areas40or pixels in the array, the physical dimensions of the array, as well as the pixel density or resolution of detector35.

Each of data readout lines48is in electrical communication with an output of a respective transistor42. For example, each of data readout lines48is associated with a row or column of transistors42, and the output (e.g., source or drain) of each transistor42in the row or column is in electrical communication with the same data readout line48such that there is one data readout line per row or column. Data readout lines48are susceptible to interference, such as electronic noise from a surrounding environment, which affects data signals being transmitted on data readout lines48. Data readout lines48are formed of a conductive material, such as a metal, and are configured to facilitate transmission of electrical signals, corresponding to incident X-rays, to image processing circuitry, for example, within dose measuring circuit30.

Scan lines50are in electrical communication with inputs (e.g., gates) of transistors42. For example, each of scan lines50are associated with a row or column of transistors42and the input of each of transistors42in the same row or column is in electrical communication with one of scan lines50. Electrical signals transmitted on scan lines50are used to control transistors42to output data on the transistor's output such that each transistor42connected to one of scans lines50are configured to output data concurrently and data from each transistor42connected to one of scan lines50flows through data readout lines48in parallel. In various embodiments, scan lines50and data readout lines48extend perpendicularly to one another to form a grid. Scan lines50are formed of a conductive material, such as a metal, and are configured to facilitate transmission of electrical signals from a controller (e.g., system controller28) to an input of transistors42.

Continuous photosensor38is deposited over transistors42, data readout lines48, and/or scan lines50. Photosensor38is formed from one or more photoelectric materials, such as one or more organic (i.e., carbon-based) and/or inorganic (i.e., non-carbon-based) materials that that convert light into electric current. In the present embodiment, the photoelectric material extends continuously as a unitary structure over the array of transistors42, data readout lines48, and scan lines50such that the photoelectric material of photosensor38substantially overlays and/or covers pixel areas40. By using a continuous unpatterned photoelectric material that is disposed over the transistor array, the density of transistors42in the array, and therefore, the pixel density of the detector, is increased as compared to patterned photosensors and/or a complexity of detector fabrication is reduced.

Electrodes, such as, but not limited to, electrical contacts of photosensor38define anode(s) and cathode(s) of photosensor38and are formed of a conductive material, such as, for example, indium tin oxide (ITO). For example, photosensor38includes electrodes disposed on a first side of photosensor38for electrically coupling the first side of photosensor38to transistors42and includes one or more electrodes disposed on a second opposing side of photosensor38for electrically coupling the second side of photosensor38to a bias voltage or vice versa. The electrodes of photosensor38form the anode(s) or cathode(s) of photosensor38.

A dielectric layer56is disposed over continuous photosensor38and a conductive layer54is disposed on dielectric layer56. Dielectric layer56includes vias58to electrically couple conductive layer54to the electrode(s) of photosensor38to allow a common bias voltage to be applied at each pixel area40of detector35.

Scintillator44is disposed over conductive layer54and generates the optical photons when exposed to X-rays. The optical photons emitted by scintillator44are detected by photosensor38, which converts the optical photons to an electrical charge that is output through transistors42to data readout lines48. Scintillator44may be a thin film deposited material or a plastic sheet with scintillator particle embedded in plastic binder, laminated using adhesive. Additionally, in various embodiments, scintillator44is directly printed onto substrate36or the photosensor38.

FIG. 3is a perspective view of an exemplary C-arm X-ray fluoroscopy system300. In the exemplary embodiment, a C-arm assembly302is used with a patient examining table assembly304to perform medical operations on a patient (not shown inFIG. 3). A large-format digital X-ray dosimeter, also referred to as a dosimetry detector306is positioned on or within a support platform308of patient examining table assembly304. Dosimetry detector306extends along a longitudinal axis310of patient examining table assembly304for a predetermined distance312. In one embodiment, distance312is equal to a length of patient examining table assembly304. In various embodiments, distance312is less than the length of patient examining table assembly304. For example, in one embodiment, distance312is less than a length of a volume of examining table below patient examining table assembly304such that dosimetry detector306and support platform308receive stream of X-ray radiation16prior to a patient (not shown inFIG. 3) receiving stream of X-ray radiation16and before a portion of stream of X-ray radiation16that has passed through the patient reaches an imaging detector314. In one embodiment, dosimetry detector306is sized to cover greater than ten per-cent of a surface area of patient examining table assembly304. In another embodiment, dosimetry detector306is sized to cover greater than fifty per-cent of a surface area of patient examining table assembly304. In some embodiments, patient examining table assembly304includes a depression315formed in the surface of patient examining table assembly304where depression315is sized complementary to a size and thickness of dosimetry detector306. Dosimetry detector306is adhesively coupled to the surface of patient examining table assembly304or is formed integrally with the surface of patient examining table assembly304. In various embodiments, patient examining table assembly304is embodied in a cardiology table, a surgical table, or an angiography table.

FIG. 4is a perspective view of C-arm X-ray fluoroscopy system300. In the exemplary embodiment, a flux field402of X-ray radiation being detected by dosimetry detector306is visually depicted on dosimetry detector306where variations in the shading of the depiction are indications of the variation in flux field intensity. The visual depiction of flux field402illustrates the difficulty of using a point dosimeter that only measures the flux at a particular point. An intensity of flux field402is shown to vary widely over a relatively small area. A small displacement of the point dosimeter from a position assumed in an estimation calculation results in a large error in the peak skin dose calculation.

FIG. 5is a perspective view of C-arm X-ray fluoroscopy system300. In the exemplary embodiment, flux field402of X-ray radiation being detected by dosimetry detector306is visually depicted on dosimetry detector306while a patient502is positioned on examining table assembly304. The visual depiction of flux field402illustrates the difficulty of using a point dosimeter that only measures the flux at a particular point. An intensity of flux field402is shown to vary widely over a relatively small area and over a portion of patient502that is subject to an examination or procedure using positioning of examination implements within patient502while observing the placement using fluoroscopy system300. A small displacement of the point dosimeter from a position assumed in an estimation calculation results in a large error in the peak skin dose calculation for the patient, the doctor, and other medical personnel assisting in the procedure.

FIG. 6is a plan view of a layout of pixels of dosimetry detector306. In the exemplary embodiment, dosimetry detector306is used for a real-time active measurement of peak skin dose to a patient during medical procedures. As used herein, real-time refers to outcomes occurring at a substantially short period after a change in the inputs affecting the outcome, for example, events occurring in real-time occur without substantial intentional delay.

Dosimetry detector306includes a substrate602on which layers of pixels604of X-ray radiation detector structure are formed as described below. In the exemplary embodiment, pixels604are formed in columns and rows spaced approximately uniformly with respect to each other. An interconnect606couples each pixel to read-out electronics using a tape-automated bonding (TAB) connection608. Each of pixels604are sized to a predetermined length l and width w. In the exemplary embodiment, l and w are selected to be approximately 2.0 millimeters (mm) In other embodiments, l and w are selected to be approximately 1 centimeter (cm). In various embodiments, l and w are selected to be unequal distances. For example, a pixel resolution of dosimetry detector306may vary across the length or width of dosimetry detector306. A particular patient examining table assembly304may be associated with a particular application, such as, but, not limited to heart applications. Because a major portion of the radiation dose is expected in the vicinity proximate the placement of the heart over a certain area of patient examining table assembly304, the pixel resolution in that area may be selected to be greater in that area (i.e., l and/or w are relatively smaller values). Because pixels604are sized so large, each of pixels604is addressed individually and multi-plexing of the signals from pixels604is not used. Also because pixels604are so large, dosimetry detector306is not particularly suitable for digital X-ray imaging. This is one differentiating factor between an imaging digital X-ray detector, such as detector35(shown inFIG. 2) and dosimetry detector306. The large size of pixels604also permits dosimetry detector306to be fabricated at relatively low cost. Accordingly, dosimetry detector306is low cost, thin and flexible, and is incorporated into an existing framework of surgical systems.

FIG. 7is a side elevation view of pixel604using a top-side interconnect.FIG. 8is a side elevation view of pixel604using a bottom-side interconnect. In the exemplary embodiment, pixel604is formed on flexible substrate602. A total thickness700of pixel604and substrate602is approximately one-eighth inch (3.175 millimeters). An anode layer702, which is formed by a conductive metal or conductive oxide and an interconnect606are printed on substrate602. A photodiode layer704is an un-patterned thin film photodiode, which, in various embodiments, is formed of plasma-enhanced chemical vapor deposition (PECVD) amorphous silicon (a-Si) or a solution coated organic photodiode (OPD). A transparent conductive metal or oxide electrode layer706is deposited over photodiode layer704. In various embodiments, photodiode layer704is sealed or encapsulated with a transparent barrier layer707, such as, but not limited to, glass to form a protective environmental seal for photodiode layer704. A scintillator layer708is then attached or deposited onto electrode layer706or barrier layer707. A thickness710of scintillator layer708is selected to be just large enough to generate a signal for a minimum detectable dose, which limits the absorption of x-rays in the detector. In the exemplary embodiment, a TFT array is not needed because an image is not being stored. Scintillator layer708may be a thin film deposited material or a plastic sheet with scintillator particle embedded in plastic binder, laminated using adhesive.

Because pixels604are relatively large, as compared to, for example, imaging detector pixels, each individual pixel has its own data channel that it feeds out information to. Consequently, because the number data channels to read out is relatively small, the data channels do not need to be multiplexed, further making the electronic read-out circuitry simpler. Pixels604are formed differently than for an imaging detector because when forming an image, it is desirable to stop or count as many X-rays as possible and convert that information into light. In the exemplary embodiment, it is desirable to form scintillator layer708just thick enough to generate an X-ray signal by stopping as few X-rays as possible to generate the signal. Scintillator layer708is tuned to have a coefficient of transmission for X-rays that is less than or equal to a coefficient of transmission for X-rays of support platform308(shown inFIGS. 3 and 5). In this way a portion of the total X-ray flux passing through support platform308and detector306(shown inFIGS. 3 and 5) is sufficient for performance of the medical imaging being done on the patient. Photodiode layer704generates electrical current which increases as a function of X-ray flux during X-ray exposure. The electrical current is conducted through metallic interconnects606, which are printed or patterned onto flexible substrate602, as shown inFIG. 7or through metallic interconnects806as shown inFIG. 8. Conductors (not shown inFIGS. 7 and 8) are either on the top (photodiode) side of substrate602(as shown inFIG. 7) and connected to an anode, or on the bottom side (opposite the photodiode) and electrically connected thru a via712in substrate602(shown inFIG. 8). At the edge of substrate602, a series of TAB bond pad connections608(shown inFIG. 6) are provided for connection to data acquisition circuit26(shown inFIG. 1), which are used to amplify and convert the current to digital information. Dosimetry detector306is incorporated into existing examining tables, such as, but, not limited to, support platform308either internally or laminated to the surface of support platform308.

FIG. 9is a side elevation view of another embodiment of pixel604using a back-side scintillator902. In the exemplary embodiment, pixel604is formed on a substrate904. A total thickness906of pixel604is approximately one-eighth inch (3.175 millimeters). An anode layer908, which is formed by a conductive metal or conductive oxide and an interconnect910are printed on substrate904. A photodiode layer912is an un-patterned thin film photodiode, which, in various embodiments, is formed of plasma-enhanced chemical vapor deposition (PECVD), amorphous silicon (a-Si), or a solution coated organic photodiode (OPD). An opaque and reflective conductive metal or oxide electrode layer914is deposited over photodiode layer912. An optional barrier layer916, such as, but not limited to, glass forms a protective environmental seal for electrode layer914. Scintillator layer902is attached or deposited onto substrate904on a side opposite to that of photodiode layer912, electrode layer914, and optional barrier layer916. A reflector layer918is deposited over scintillator layer902. A thickness710of scintillator layer708is selected to be just large enough to generate a signal from photodiode layer912for a minimum detectable dose, which limits the absorption of x-rays in the detector.

Embodiments of this disclosure describe incorporating a peak skin dose measurement device or dosimeter within the examining table that yields peak skin dose accumulated as a function of position, with a size resolution of approximately 2-3 millimeter (mm) to approximately 1 centimeter (cm). The dosimeter is substantially transparent to the x-ray radiation, stopping only a few percent of the delivered dose. The dosimeter is placed in close proximity to the skin of a patient to accurately measure effects, such as, scattering which is influenced by patient placement, size and examining table geometry. The dose measurement device or dosimeter is part of an integrated dose detection system that works by placing a low cost, large area array of high x-ray transparency detectors onto a thin, flexible substrate. The x-ray detector pixels are formed to be 2.0-3.0 mm in size and 1.0 cm or larger.

It will be appreciated that the above embodiments that have been described in particular detail are merely example or possible embodiments, and that there are many other combinations, additions, or alternatives that are included.

Based on the foregoing specification, the above-discussed embodiments of the disclosure are implemented using computer programming or engineering techniques including computer software, firmware, hardware or any combination or subset thereof. Any such resulting program, having computer-readable and/or computer-executable instructions, are embodied or provided within one or more computer-readable media, thereby making a computer program product, i.e., an article of manufacture, according to the discussed embodiments of the disclosure. The computer readable media may be, for instance, a fixed (hard) drive, diskette, optical disk, magnetic tape, semiconductor memory such as read-only memory (ROM) or flash memory, etc., or any transmitting/receiving medium such as the Internet or other communication network or link. The article of manufacture containing the computer code is made and/or is used by executing the instructions directly from one medium, by copying the code from one medium to another medium, or by transmitting the code over a network.

While the disclosure has been described in terms of various specific embodiments, it will be recognized that the disclosure may be practiced with modification within the spirit and scope of the claims.

The above-described embodiments of a system and method of radiation dosimetry provides a cost-effective and reliable means for measuring a peak skin dose of radiation delivered to a patient and/or medical professional during a procedure where radiological imaging is also performed. As a result, the system and method described herein facilitate monitoring and reducing dose to patients and medical professionals in a cost-effective and reliable manner.