Patent Description:
This application claims priority to U. No.<CIT>, and entitled A PROTON IMAGING SYSTEM FOR OPTIMIZATION OF PROTON THERAPY.

The Bragg peak phenomena enables particles such as protons or ions to precisely target tumors for radiation therapy, while healthy tissues receive a minimal dose when compared with x-ray therapy systems. Proton radiation therapy however requires precise patient alignment, and also adjustment of initial proton energy so that the maximum dose corresponding to the Bragg peak is deposited in intended tissues. In order to adjust the range of a proton beam, so that the maximum dose corresponding to the Bragg peak is deposited in intended tissues, treatment planning may require three-dimensional map of a particular patient in terms of relative stopping power, or the energy loss of the proton beam in a material relative to that of water. <CIT> discloses a proton computed tomography (pCT) detector system, including two tracking detectors in sequence on a first side of an object to be imaged, two tracking detectors in sequence on an opposite side of the object to be imaged, a calorimeter, and a computer cluster, wherein the tracking detectors include plastic scintillation fibers. A method of imaging an object comprises emitting protons from a source through two tracking detectors, through and around the object, and through two opposite tracking detectors, detecting energy of the protons with a calorimeter, and imaging the object. <NPL>, "Uzunyan", discloses a proton CT scanner. <CIT> relates to charged-particle segmented strip detectors. P Pemler et. "A detector system for proton radiography on the gantry of the Paul-Scherrer-Institute", discloses a proton radiography system with a magnetic deflection sweeper magnet for deflecting a scanning proton pencil beam.

The present disclosure relates to a method for operating a medical imaging system. The method comprises:.

Embodiments of a medical imaging system, such as a proton radiography system or a proton tomography system, are described. In the presently claimed invention, a proton radiography system includes a first tracking detector, a second tracking detector, and a residual range detector. In this embodiment, the first and second tracking detectors use a fiber bundling architecture that, among other things, significantly improves proton transverse position resolution while reducing the complexity of the proton radiography system. The residual range detector uses a light collection architecture that, among other things significantly improves proton residual range resolution while simplifying the overall residual range detector design.

Some variations of the fiber bundling architecture used in the first and second tracking detectors utilize a planar substrate and include at least two layers of scintillating fibers on each side of the substrate that extend along particular directions. More specifically, each one of the fibers on one side of the substrate extends along a direction that is perpendicular to a direction of each one of the fibers on the opposite side of the substrate.

Some variations of the bundling architecture also use (a) an adjacent layer bundling scheme and/or (b) a lateral strip-based bundling scheme. In the adjacent layer bundling scheme, one fiber from each layer on one side of the substrate are bundled together to form a fiber doublet. In the lateral strip-based bundling scheme, groups of fibers or fiber doublets on each side of the substrate are segmented or partitioned into strips or sections of a particular number of fiber or fiber doublets (e.g., a strip of ten fibers or ten fiber doublets). In some examples, particular fibers or fiber doublets (e.g., the fourth fiber or the fourth fiber doublet) of each strip are bundled together and coupled to a same anode of a particular multi-anode light detector (or other type of photon detector).

Many benefits and advantages flow from the fiber bundling architecture of the present technology. For example, the number of light detectors and electronics channels required to resolve proton transverse position are substantially reduced, while at the same time proton transverse position resolution approaches <NUM> or better, when compared to conventional proton radiography systems. By extension, the described fiber bundling architecture addresses many of the issues associated with conventional proton radiography systems, which are also bulky and expensive both in terms of component and data processing costs.

Regarding the light collection architecture, the residual range detector includes a scintillator material that, in some examples, is a bulk piece of material. The residual range detector also includes at least one light detector coupled to the scintillator material at a surface of the scintillator material. If the at least one light detector comprises a plurality of light detectors, they may be arranged to exhibit a unit cell pattern selected from, for example, square, rectangular and polygonal. Additionally, in some examples, an anti-reflective material (e.g., a thin film, other coating, or cover) is present on at least one surface (or multiple surfaces) of the scintillator material.

Many benefits and advantages flow from the light collection architecture of the present disclosure. For example, both the amount of time required to collect scintillation light and scintillation light collection efficiency are substantially improved, while at the same time residual range resolution approaches <NUM> or better per proton, when compared to conventional proton radiography systems. By extension, the disclosed light collection architecture addresses many of the issues associated with typical proton radiography systems, which also suffer from proton residual range resolution error that propagates through to map reconstruction during the treatment planning process.

Exemplary embodiments of a medical imaging system that utilize one or more of the above features are described in further detail below. While protons are used an example particle for use with the system described below, other particles (e.g., ions of other elements, including helium ions, lithium ions, beryllium ions, carbon ions, boron ions, deuterons) could also be used. Additionally, while a proton radiograph system is described below, the present technology could also be used in other medical imaging systems, such as a proton tomography system.

<FIG> depicts system <NUM>, which is an embodiment of a proton radiography system according to the present technology. In some examples, a pencil beam <NUM> (alternatively, a broad beam could also be used with variations of the present technology) of protons (although other particles, such as heavy ions, could also be used) is generated or extracted from a source (see <FIG>) and scanned across a field by a scanning element (see <FIG>). For example, pencil beam <NUM> is scanned across a region of interest of object <NUM>, which is a human head in <FIG>. The position of pencil beam <NUM> as it enters object <NUM> can be determined based on data from tracking detector <NUM> that generates photons at locations where protons traverse tracking detector <NUM>, as described in more detail below. Similarly, the position of pencil beam <NUM> as it exits object <NUM> can be determined using data from tracking detector <NUM>. Note that tracking detectors <NUM> and <NUM> are spaced apart to allow object <NUM> to be positioned between them. In some embodiments, only one tracking detector is needed or two or more tracking detectors are used on one or both sides of object <NUM>. Potential architectures for tracking detectors <NUM> and <NUM> are described in more detail below.

Residual range detector <NUM> is positioned adjacent tracking detector <NUM>. Residual range detector <NUM> includes scintillator material <NUM> (represented in <FIG> by a box). In one example, the scintillator material may be one sold by Eljen Corporation. As a proton of pencil beam <NUM> enters scintillator material <NUM> through the surface facing tracking detector <NUM>, the proton generates photons <NUM> (represented by dotted lines in <FIG>) as the protons loses energy from interacting with scintillator material <NUM>. These photons can then be collected by photon detectors (<FIG> and <FIG>) coupled to scintillator material <NUM> on the surface of scintillator material <NUM> that is opposite the surface facing tracking detector <NUM>. The coupling of the photon detectors is depicted in <FIG> by circles <NUM>. The signal generated by the photon detectors is proportional to a residual energy of a proton as it entered scintillator material <NUM>. This information combined with the initial energy of the proton and the location of the proton as it entered and exited object <NUM>, along with similar information for many additional protons, can be used to generate image <NUM> of object <NUM>. By using multiple proton energies and/or protons at different angles (e.g., <NUM> different angles) in a proton tomography system using the present technology, 3D images can also be produced.

Note that reference axes <NUM> show that pencil beam <NUM> is traveling along the z-axis and tracking detectors <NUM> and <NUM> are perpendicular to the z-axis. <FIG> and <FIG> described below will be described with respect to the same reference axes.

<FIG> depicts system <NUM> of <FIG> with additional detail and from a different perspective, as indicated by reference axes <NUM>. With reference to <FIG>, system <NUM> includes source <NUM> that generates or extracts pencil beam <NUM>. Scanning element <NUM> is used to steer pencil beam <NUM> and includes scanning magnets 204a and 204b. Scanning magnets 204a scan the pencil beam <NUM> in x-directions as defined with respect to reference axes <NUM>, and scanning magnets 204b (the second of which is hidden in the perspective of <FIG>) scan pencil beam <NUM> in y-directions as defined with respect to respect to reference axes <NUM>. In general, scanning element <NUM> is programmable such that pencil beam <NUM> is scan-able across the entirety of the field in any pattern, and, as determined by source <NUM>, at any initial energy at any point in time. In the presently claimed invention, source <NUM> is capable of controlling or varying the initial energy of the protons of pencil beam <NUM>. During a scan, the extent of the field is in general limited by the planar dimensions of first tracking detector <NUM>, second tracking detector <NUM>, or residual range detector <NUM> of system <NUM>. Example field areas include 10x10 cm<NUM> to <NUM>. <NUM><NUM>.

By using in the invention a different initial energy at different transverse positions of the tracking detectors and the object being imaged, the depth of the residual range detector can be kept small. For example, the initial energies can be chosen to keep the residual range between <NUM> and <NUM> across the field to be imaged regardless of the thickness or density of the object along a particular path. The more range in initial energy that is possible, smaller residual range detectors may be possible.

In some examples of an architecture for tracking detector <NUM>, as protons of pencil beam <NUM> traverse first tracking detector <NUM>, the protons interact with fibers either side of substrate <NUM>. Specifically, each side of substrate <NUM> includes, for example, two layers of fibers, i.e., fibers <NUM> on one side and fibers <NUM> on the opposite side of substrate <NUM>. Fibers <NUM> and <NUM> may be scintillating fibers so that when a proton impinges a fiber, the scintillating properties of the fiber will cause one or more photons to be generated. These photons are captured by light detector <NUM>, which generates an electrical signal based on the detected photons. The electrical signal is transmitted to computing system <NUM>. By knowing the location and orientation of fibers <NUM> and <NUM> that produced photons, the location of the proton that traversed tracking detector <NUM> may be determined by computing system <NUM>. Additionally, computing system <NUM> may also use data from scanning element <NUM> to determine or verify the location where a proton passed through tracking detector <NUM>. Once the location is known the initial directional vector of pencil beam <NUM> can also be determined based on the focal point of source <NUM>.

In some examples, fibers <NUM> are oriented perpendicular to fibers <NUM>. If light detector <NUM> indicates that a proton passed through one fiber of fibers <NUM> and one fiber of fibers <NUM> and computing system <NUM> knows the location of these two fibers, computing system <NUM> can determine that the X-Y coordinate on tracking detector <NUM> where the proton traversed tracking detector <NUM> is at the intersection of the two fibers. Additionally, if fibers of fibers <NUM> or fibers of fibers <NUM> are connected together to reduce the number of detectors needed in light detector <NUM>, then the estimated expected position of pencil beam <NUM> based on data from or instructions sent to scanning element <NUM> may be used in determining the X-Y coordinate where the proton traversed tracking detector <NUM>, as described in more detail below with respect to <FIG>.

As protons of pencil beam <NUM> traverse object <NUM>, protons may be scattered, as is depicted in <FIG> as an exaggerated change in direction of pencil beam <NUM> in object <NUM>. After the protons exit object <NUM>, the exit location where protons traverse tracking detector <NUM> can be determined in a similar manner as described above with respect to tracking detector <NUM>. Similar to tracking detector <NUM>, tracking detector <NUM> includes fibers on substrate <NUM>. Specifically, each side of substrate <NUM> includes, for example, two layers of fibers, i.e., fibers <NUM> on one side and fibers <NUM> on the other side of substrate <NUM>. Fibers <NUM> and <NUM> may be scintillating fibers so that when a proton impinges a fiber, the scintillating properties of the fiber will cause one or more photons to be generated. These photons may be captured by light detector <NUM>, which may generate an electrical signal based on the detected photons. The electrical signal is transmitted to computing system <NUM>. By knowing the location and orientation of fibers <NUM> and <NUM> that produced photons, the location of the proton that traversed first tracking detector <NUM> may be determined by computing system <NUM>. Additionally, computing system <NUM> may also use information from scanning element <NUM> to determine or verify the location on tracking detector <NUM> where a proton passed through tracking detector <NUM>.

In some examples, fibers <NUM> are oriented perpendicular to fibers <NUM>. If light detector <NUM> indicates that a proton passed through one fiber of fibers <NUM> and one fiber of fibers <NUM> and computing system <NUM> knows the location of these two fibers, computing system <NUM> can determine that the X-Y coordinate on tracking detector <NUM> where the proton traversed tracking detector <NUM> is at the intersection of the two fibers. Additionally, if fibers of fibers <NUM> or fibers of fibers <NUM> are connected together to reduce the number of detectors needed in light detector <NUM>, then the estimated or expected position of pencil beam <NUM> based on data from or instruction sent to scanning element <NUM> may be used in determining the X-Y coordinate on tracking detector <NUM> where the proton traversed tracking detector <NUM>, as described in more detail below with respect to <FIG>.

While an exemplary architecture of a tracking detector has been described, other architectures are possible. For example, if fibers are rigid enough, the fibers could be bonded together to avoid using a substrate. As another example, with respect to tracking detector <NUM>, fibers <NUM> and <NUM> could be placed on the same side of substrate <NUM> or fibers <NUM> and <NUM> could be placed on separate substrates that are placed next to each other.

As protons of pencil beam <NUM> enter residual range detector <NUM>, they impinge scintillator material <NUM> and generate photons that are collected by photon detectors <NUM> (while four photon detectors are depicted in <FIG>, system <NUM> includes sixteen photon detectors as indicated by the circles representing the photon detectors couplings to scintillator material <NUM> in <FIG>). Photon detectors <NUM> are, for example, photomultiplier tubes or other similar devices. Photon detectors <NUM> generate electrical signals based on the number of photons collected and generate electrical signals that are provided to computing system <NUM>, which may calculate values such as total energy. Based on the electrical signals, and potentially other information (such as the X-Y coordinate of where a proton exited object <NUM> and traversed tracking detector <NUM>), computing system <NUM> may determine a residual energy for a proton of pencil beam <NUM> that entered scintillator material <NUM> after exiting object <NUM>.

Protons enter scintillator material <NUM> via surface <NUM>. Generated photons are collected by photon detectors <NUM> as the photon exit surface <NUM> of scintillator material <NUM>. The dimensions of scintillator material <NUM> may be selected to ensure that protons stop in scintillator material <NUM> as opposed to passing through scintillator material <NUM>. This ensures that protons of pencil beam <NUM> generate a large number of scintillation photons within a few nanoseconds. Surface <NUM>, surface <NUM>, surface <NUM>, and/or the other two surfaces of scintillator material <NUM> not depicted in <FIG> are, in some examples, covered (e.g., deposited, coated, or arranged next) with an anti-reflective or photon absorbing material. For example, the walls of scintillator material <NUM> are painted black. The anti-reflective material ensures that mainly direct photons that have not scattered off the walls of scintillator material <NUM> are collected at photon detectors <NUM>. The anti-reflective material may include different materials on different surfaces of scintillator material <NUM>. In one example, the anti-reflective material may be Eljen Corporation EJ510B black paint. The anti-reflective material may absorb <NUM>% or more of the photons that contact the material. The anti-reflective material adds to the high speed operation of system <NUM>.

The use of multiple photon detectors also provides the potential to obtain additional position data for the location that a proton exited object <NUM>. For example, with reference to <FIG>, if photon detectors are coupled to scintillator materials <NUM> as indicated by circles <NUM> (sixteen total), the photon detector nearest where the photon entered scintillator material <NUM> should produce the strongest signal. If the position of the photon detector that produces the strongest signal does not correlate with the position indicated by the signals generated from tracking detector <NUM> and light detector <NUM>, then an event that should be rejected may exist, such as inelastic scatter.

<FIG> depicts two cross-sections of a tracking detector that may be used to implement tracking detectors <NUM> and <NUM> (<FIG> and <FIG>). Cross-section <NUM> depicts tracking detector along the plane parallel to the z-axis and x-axis, as depicted in reference axes <NUM>. Cross-section <NUM> depicts the same tracking detector along the plane parallel to the y-axis and x-axis, as depicted in reference axes <NUM>, which is also perpendicular to the plane of cross-section <NUM>. Additionally, cross-section <NUM> is zoomed-out as compared to cross-section <NUM>. In cross-section <NUM>, the direction of protons is along the z-axis as depicted by path <NUM>. In cross-section <NUM>, the direction of protons is coming out of the figure.

As depicted in cross-section <NUM>, the tracking detector includes a substrate <NUM> having two layers of fibers, <NUM> and <NUM>, respectively, on one side and two layers of fibers, <NUM> and <NUM>, respectively on the other side. Layers of fibers <NUM> and <NUM> are laid out perpendicular to layers of fibers <NUM> and <NUM>. Only one fiber of layer of fibers <NUM> and one fiber of layer of fibers <NUM> are visible because the other fibers are blocked from view. Layer of fibers <NUM> includes fibers 312a-<NUM> and layer of fibers <NUM> includes fibers 314a-<NUM>. As depicted in <FIG>, the layers of fibers on each side of the substrate may be offset from each other so that fibers of one layer (e.g., layer <NUM>) are positioned between two fibers in the adjacent layer (e.g., layer <NUM>). In other words, the fibers in one layer may be offset from the other layer by about one half the width of a single fiber. In this layout, protons that go through the interface between two fibers in one layer, should also go through the middle of the fiber in the next layer, which results in higher efficiency. Other architectures could have additional layers of fibers or only a single layer of fibers.

Fibers of adjacent layers can be bundled together so they connect to a single light detector channel. For example, with reference to <FIG>, fibers 312a and 314a can be bundled together into fiber doublet <NUM> so fibers 312a and 314a connect to a single channel of the light detector. In some cases, bundling occurs by combining the ends of fibers 312a and 314a in parallel so that the outputs of fibers 312a and 314a can be detected together.

Multiple fibers or fiber doublets may be organized in logical strips. For example, if fibers 312a-312d are respectively bundled with fibers 314a-314d to form four fiber doublets, the four fiber doublets may be treated as strip <NUM>. Similarly, if fibers 312e-<NUM> are respectively bundled with fibers 314e-<NUM> to form four fiber doublets, the four fiber doublets may be treated as strip <NUM>. A strip may include more fibers or fiber doublets, such as <NUM> fibers or fiber doublets.

To further reduce the number of channels required in a light detector, similarly positioned fibers or fiber doublets in strips on a side of substrate <NUM> may be bundled together and connected to a single channel of the light detector. In this case, the location of the fiber or fiber doublet that generated photons within the strips in combination with the expected location of the pencil beam can be used to located the position of the pencil beam accurately to within, for example, <NUM> when using <NUM><NUM> fibers. For example, the expected location of the pencil beam can be used to determine the expect strip where the proton will be and the fiber or fiber doublet that produces photons can be used to identify the location within the strip of the proton.

The two types of bundling described above (i.e., bundling adjacent fibers of different layers and bundling fibers or fiber doublets of different strips) can be used together or separately in different variations of the present technology.

Cross-section <NUM> depicts strips <NUM>, <NUM>, <NUM>, and <NUM>, which of which includes a plurality of fiber doublets. For example, strip <NUM> includes four fiber doublets made of fibers 312a-312d and fibers 314a-314d (e.g., fiber doublet <NUM>), as described above. Similarly, strip <NUM> include fiber doublet <NUM>, strip <NUM> include fiber doublet <NUM>, strip <NUM> includes fiber doublet <NUM>, and strip <NUM> includes fiver doublet <NUM>. Because fiber doublets <NUM>, <NUM>, <NUM>, and <NUM> are located at the same positions within strips <NUM>, <NUM>, <NUM>, and <NUM>, the outputs of these fiber doublets may be bundled together and connected to light detector <NUM> via a single channel (i.e., channel <NUM>). Light detector <NUM> includes several channels (not shown) that are each connected to a group of fiber doublets that have the same position in the strips of the tracking detector. Light detector <NUM> provides electrical signals to computing system <NUM> representative of photons generated in the fibers of the tracking detector as received by the channels (e.g., channel <NUM>) of light detector <NUM>.

While an exemplary architecture of a tracking detector has been described in <FIG>, the same bundling architecture could be used in other configurations of tracking detectors. For example, with respect to the tracking detector of <FIG>, if fibers are rigid enough, fiber layers <NUM>, <NUM>, <NUM>, and <NUM> could be bonded together to avoid using a substrate. As another example, fiber layers <NUM> and <NUM> could be placed on the same side of substrate <NUM> as fiber layers <NUM> and <NUM>. In another example, fiber layers <NUM> and <NUM> could be placed on a separate substrate than fiber layers <NUM> and <NUM> and the two substrates could be placed next to each other.

<FIG> illustrates a proton bunch train <NUM> whereby individual protons 402a-d are superimposed on a sinusoid <NUM> that represents an RF (Radio Frequency) accelerating field. In some examples, bunches with more than one proton (e.g., protons 402c-d) are rejected for image analysis or map reconstruction but will still contribute to radiation dose. Bunches in adjacent cycles (e.g., protons 402a-b) may be used for image analysis or map reconstruction if the residual range detector recovers in time to detect protons that impinge in immediately subsequent cycles.

<FIG> depicts residual range detector <NUM> that could be used to implement residual range detector <NUM> of system <NUM> of <FIG>. Residual range detector <NUM> includes scintillator material <NUM>, which, in some examples, is a bulk scintillator material. Scintillator material <NUM> may be contained within an enclosure <NUM> that provides for an anti-reflective surface for photons generated in scintillator material <NUM>. Alternatively, scintillator material <NUM> may also have an anti-reflective coating (e.g., a thin film) deposited on the surfaces (e.g., surface <NUM>) of scintillator material <NUM>. Photons <NUM> are generated in scintillator material <NUM> by a proton entering scintillator material <NUM> along path <NUM>. Only direct photons are collected by the photon detectors <NUM>, this enhances light collection efficient and speed.

<FIG> illustrates multiple unit cell patterns of photon detectors coupled to the scintillator material <NUM> of <FIG>. The coupling locations of the photon detectors are indicated by circles <NUM>. Pattern in example <NUM> is a square unit cell pattern <NUM>. Pattern in example <NUM> is a rectangular unit cell pattern <NUM>. Pattern in example <NUM> is a polygonal unit cell pattern <NUM>. Pattern in example <NUM> is a square pattern <NUM> where only a single detector is present. Trade-offs govern which pattern is selected, e.g., polygonal unit cell pattern <NUM> might be selected if very high precision is required and cost and bulkiness is not an issue.

<FIG> depicts residual range detector <NUM> and simulated generated photons <NUM> that are generated by a proton interacting with the scintillator material. Circles <NUM> depict the locations of photon detectors coupled to the scintillator material. The photon detectors capture a certain number of generated photons <NUM> and produced a corresponding electrical signal representative of the number of photons that were captured. The collection efficiency of residual range detector <NUM> varies based on the location that the proton entered the scintillator material. Lower collection efficiency may result in higher uncertainty in the residual range of a proton. For example, with respect to the uncertainty data in graph <NUM>, which represents data for a quadrant of residual range detector <NUM>, the collection efficiency at the edge of the scintillator material (further from the center along the x or y axis) is lower because the sides of scintillator material are anti-reflective. This results in a higher uncertainty for the residual range near the edges of residual range detector <NUM>. Graph <NUM> also shows that higher uncertain may occur when a proton enters scintillator material in between where the photon detectors are coupled to the scintillator material.

<FIG> depicts residual range detector <NUM> and simulated generated photons <NUM> with a different configuration of photon detectors (not shown). As can be seen by the locations of circles <NUM> that represent the locations where photon detectors are coupled to the scintillator material, the photon detectors in residual range detector <NUM> are spaced closer together than in residual range detector <NUM> of <FIG>. As can be seen in the residual range uncertainty in graph <NUM>, the closely spaced photon detectors in residual range detector <NUM> can determine the residual range of protons that enter the center of residual range detector <NUM> with less and more uniform uncertainty as compared to residual range detector <NUM> of <FIG>. The uncertainty at the edge of residual range detector <NUM>, however, is higher than the uncertainty at the edge of residual range detector <NUM> of <FIG>.

The differences in performance between residual range detectors <NUM> and <NUM> of <FIG> and <FIG>, respectively, show that a residual range detector can be configured according to advantageous specifications. For high accuracy and large detection fields, a large number of closely spaced photon detectors may be needed. If, however, slightly lower accuracy is acceptable, a lower number of more widely spaced photon detectors could be used to reduce cost and complexity.

<FIG> shows a detector gain calibration system <NUM> that, in some examples, is used to calibrate a residual range detector (such as residual range detector <NUM> of system <NUM> in <FIG> and <FIG>) having scintillator material <NUM> and photon detectors <NUM>. Calibration system also includes pulse generator <NUM>, UV LED <NUM>, photodiode <NUM>, and diffuser <NUM>. In other example calibration systems, a diffuser may not be used.

In one example, in a residual range detector that uses a 4x4 array of photon detectors, such as photomultiplier tubes (PMT), each PMT is labeled PMT_ij where i and j and the row and column of each PMT. Using analog summing electronics, we form three output signals from sixteen input signals. This allows three channels to be digitized channels instead of sixteen, which is a major savings in cost of electronics and data volume. The three output signals are a total energy signal E, and two position dependent signals for diagonal coordinates (rather than row-column coordinates).

The procedure to reconstruct proton residual range involves a two-step process:.

The first step is to set and maintain the gains of the individual PMTs relative to a photodiode with LED pulsing, as shown in <FIG>. A photodiode provides a convenient reference with stable unity gain, as long as care is taken to ensure temperature and gain stability of any preamplifiers. In some cases the PMTs gains are within approximately <NUM>% of each other. This may cause a slight increase in the statistical error from the number of photons detected. The gains of the PMTs may be set and maintained individually to a value fixed relative to the photodiode.

Because the light collection efficiency of residual range detector <NUM> is position dependent (see discussion with respect to <FIG> and <FIG>), E may be corrected for the X and Y position of the proton track in scintillator material <NUM>. Furthermore, there are fluctuations and correlations between E, U, and V for proton tracks at a particular position, and an optimal reconstruction may take these into account.

The second step is to acquire a calibration data set, organized with protons in a 3D grid: X, Y, and residual range R. In some examples, a grid spacing of <NUM> for all <NUM> coordinates is used. X and Y can be determined from the tracking system and events within approximately <NUM> from the center of each grid point may be selected. For each grid point, labeled by X, Y, and R, the average E, U, and V, as well as the 3D covariance matrix (optionally, C can be included for a 4D covariance matrix) may be saved as calibration data for use later.

For an individual proton with measured X, Y, E, U, V, the grid of data can be used to reconstruct the residual range. For a hypothesized residual range, the following steps may be used:.

<FIG> depicts experimental results in graph <NUM> for a residual range detector according to the present technology (e.g., a residual range detector similar to residual range detector <NUM> of <FIG> and <FIG>). The pulse height measured in terms of mV shows an approximately linearly correlation with residual range of protons. The residual range detector that produced the results in graph <NUM> included a 10x10x10 cm<NUM> active volume of plastic scintillator and a 2x2 array of PMTs.

<FIG> depicts graph <NUM> containing data for spreads of pulse heights at one range setting of protons in the same experiment that produced the data of <FIG>. Curves <NUM> are signals from one PMT. Curves <NUM> are the total energy signal.

<FIG> depicts graph <NUM> of data for a tracking plane or detector simulation. Graph <NUM> shows that using two tracking planes (tracking detector <NUM> would be considered "one" tracking plane) is better than using one, because transverse positon uncertainty is lower (resolution is better) when two tracking detectors are used. Curve <NUM> is for one tracking planes positioned <NUM> after water. Curve <NUM> is for two tracking planes positioned <NUM> after water. Curve <NUM> is for one tracking planes positioned <NUM> after water. Curve <NUM> is for two tracking planes positioned <NUM> after water. As can be seen from graph <NUM>, there may be some benefit to using two tracking planes as opposed to one tracking plane. With respect to system <NUM> of <FIG>, using two tracking planes may correspond, for example, to using two tracking detectors on one or both sides of object <NUM>. In some cases when a pencil beam is used, only one tracking plane on the upstream side is necessary because the input direction is known as a function of detected proton position
<FIG> depicts an example support structure <NUM> that can be used to mount some components of system <NUM> of <FIG> and <FIG>. For example, support structure <NUM> includes base <NUM> and track <NUM>. Tracking detector <NUM>, which may correspond to tracking detector <NUM> of <FIG> and <FIG>, may be mounted on track <NUM> so that tracking detector <NUM> is moveable along base <NUM>. Tracking and residual range detector <NUM> includes a tracking detector (such as tracking detector <NUM> of <FIG> and <FIG>) and a residual range detector (such as residual range detector <NUM> of <FIG> and <FIG>). Like tracking detector <NUM>, tracking and residual range detector <NUM> is mounted on track <NUM> so that tracking and residual range detector <NUM> is moveable along base <NUM>. Tracking and residual range detector <NUM> could also be separated into two distinct components.

<FIG> depicts mounting system <NUM> for support structure <NUM>. Mounting system <NUM> includes a mounting arm <NUM> and a mounting base <NUM>. In some examples, mounting arm <NUM> and/or mounting base <NUM> is adjustable or controllable to position support structure <NUM> into different positions as is helpful for treatment or imaging of an object between tracking detector <NUM> and tracking and residual range detector <NUM>. Other example mounting systems can include the capability to rotate the angle at which an object is being imaged. With these system, some embodiments of the present technology can be used to implement tomography imaging systems that produce 3D data about the object.

Some embodiments of the disclosed detector of the present technology can transform the practice of proton radiation therapy because the same will enable facilities to efficiently and confidently deliver optimal treatments fully realizing the promise of the Bragg peak. As mentioned above, the Bragg peak enables particles such as protons or so-called heavy ions to precisely target tumors for radiation therapy while exposing healthy tissues to a smaller dose when compared with conventional radiation therapy. However, proton radiation therapy requires precise patient alignment and adjustment of proton initial energy, so that the proton beam has the correct range in the patient to stop in the tumor. In order to adjust the range of the proton beam, treatment planning uses a three-dimensional map of a patient in terms of relative stopping power, or the energy loss of protons of the beam in a material relative to that of water. Typically, such maps are obtained from x-ray computerized tomography scans; however, there are uncertainties or errors in converting x-ray absorption units (Hounsfield) to relative stopping power, and patient inhomogeneities introduce additional uncertainties. Furthermore, metallic implants or other high-density materials can cause shadowing artifacts and streaking.

Proton technology is not yet mature, and uncertainty margins are often greater than that for modern photon therapy. Thus, it is contemplated that there is a clear need to reduce alignment uncertainties, enable more complex treatments using more proton directions, deliver a higher dose to a tumor per treatment, and reduce range uncertainties such as to within <NUM> range precision. It is contemplated that such improvements, which use frequent imaging, should be accompanied by improvements in patient throughput to improve the cost-effectiveness of proton therapy relative to conventional radiation therapy. The significant expense of building and maintaining proton therapy facilities has been recognized as a major disadvantage, which has only recently been reduced by the development of more compact systems and single-room options. It is contemplated that new imaging technologies should integrate seamlessly into proton therapy systems to streamline and simplify patient operations, rather than add complexity and expense. The features or aspects of embodiments of the present technology may provide such and more benefits.

For example, the embodiments of the detector of the present technology may leverage at least one tracking detector to measure the transverse position of individual protons before and/or after a patient, and a residual range detector to determine the proton energy absorbed within the patient. Proton radiography produces two-dimensional images with a single projection angle, directly quantifying proton range through the patient rather than x-ray absorption power. Digitally reconstructed radiographs can be derived from previous x-ray computerized tomography scans and compared to proton radiographs to validate and improve relative stopping power maps and patient alignment. Proton computed tomography, by contrast, measures the three-dimensional relative stopping power map of a patient by acquiring many proton histories from many projection angles and applying advanced iterative reconstruction algorithms. While proton computed tomography may leverage the same or similar detector technology as proton radiography, proton computed tomography typically has greater operational complexity, produces larger data volumes, and uses proton energies high enough to traverse a patient in all directions. Additionally, the embodiment of the detector of the present technology may be leveraged to commercialize a proton radiography system producing two-dimensional images using protons with enough energy to traverse the patient. A subsequent radiation treatment uses a lower energy, higher intensity beam which terminates in a tumor. The use of a proton beam for both imaging and treatment streamlines patient setup and quality assurance procedures, reduces alignment uncertainties, and reduces range uncertainties.

Currently, radiation therapy is needed for more than <NUM>% of the <NUM> million Americans who are annually diagnosed with cancer. A conservative estimate from the Mayo Clinic is that <NUM>,<NUM> new cancer patients each year in the United States could benefit from proton therapy, well above current capacity. Proton radiation therapy can potentially spare large amounts of normal tissue from low to intermediate radiation dose and avoid organs at risk. This reduces late effects and improves quality of life, and is especially important for patients with high cure rates and long survival times. A policy statement issued by the American Society of Therapeutic Radiation Oncology cites scientific evidence confirming that proton beam therapy is particularly useful in a number of pediatric patients, particularly those with brain tumors, as well as for certain adult cancers requiring high doses in close proximity to critical structures. Additional research on more common cancer disease sites, such as breast, prostate and lung, is ongoing, with clinical trials accruing patients in all three disease sites from proton therapy facilities in the United States. Currently, sixteen (<NUM>) proton therapy facilities with a total of fifty-six (<NUM>) treatment rooms are operating in the United States, with many more under development. The embodiment detector of the present technology has the potential to be adopted for routine use in all treatment rooms using pencil beam scanning, which is quickly becoming the standard, rather than broad beams individually tailored for each patient.

Some embodiments of the detector of the present technology are non-complex, lightweight, easily scaled to large field sizes, operates at high speed to avoid bottlenecks in patient throughput and exposes the patient to the minimum possible dose for a given resolution. Proton radiography is not currently used routinely due to designs that are bulky, expensive and difficult to incorporate into the clinical environment. It is contemplated that an alternative is to use a single detector plane behind the patient and vary the proton energy to find the range through the patient. However, drawbacks to this approach include for example: inefficient use of proton dose since most protons do not contribute to measurement; poor spatial resolution since protons are not tracked; and, by using a broad beam for imaging, does not use the same beam system as for treatment in the case of pencil beam scanning systems.

Some embodiments of the detector of the present technology leverage fast-scintillation detector technology and the high-performance design is non-complex and monolithic, thereby reducing construction costs. For example, a low channel count is leveraged to minimize electronics development costs, and residual range resolution of <NUM> per proton or better is achieved. At this level, the range resolution is dominated by intrinsic range fluctuations rather than detector measurements for typical patient dimensions. This is helpful in order to achieve the favorable dose performance of proton radiography, which averages measurements from many protons to form an image. Doses as low as <NUM>µGy (microgray) are possible for a resolution in Water Equivalent Path Length (WEPL) of <NUM> for each square pixel.

Some embodiments of the detector of the present technology enable measurement of up to <NUM> million protons per second, resolving individual protons as close as <NUM> nanoseconds. Accelerator systems deliver proton bunches at the frequency of their RF cavities, as illustrated in <FIG>, with RF frequencies as high as <NUM>. In the case of low-intensity beams for imaging, most bunches will be empty, and the remaining bunches will contain a single proton. A <NUM> proton beam will have protons separated by an average of <NUM> nanoseconds, with a random separation distribution as close as <NUM> nanoseconds in time. As noted above in connection with <FIG>, a small fraction of bunches will have two or more protons. These may be rejected for analyses but will still contribute to the dose. Advantageously, such speeds allow for, as an example, a 20x20 cm<NUM> field to be imaged with less than a second of beam time with a resolution in WEPL of <NUM> for each square pixel.

Some embodiments of the detector of the present technology are especially advantageous for pencil beam scanning systems. Pencil beams lead to protons sequentially hitting the same region of a detector, adding to the challenge of achieving high event rate capability. Further, use of pencil beams for both imaging and treatment will be very powerful for alignment and quality assurance, compared to imaging with a broad beam for example. Further, pencil beam position setting information can add redundancy to the position reconstruction, which is helpful for rejecting events with nuclear scatters or other problems. Further, a pencil beam scanning system can be used to divide the field for a proton radiograph into regions with different proton energy settings for each region based on the estimated range in that region, as obtained from a previous x-ray computerized tomography scan. This allows the system to maintain a low residual range for the protons as the beam scans across the patient, and has several benefits including for example: the residual range detector can be thinner, such as <NUM> or less, saving on weight and volume in the treatment area, and making the read-out easier; the lower total range for the protons is more optimal for range resolution relative to dose, lower range also results in fewer protons lost to nuclear interactions, which also results in lower dose for a given image quality.

Some embodiments of the detector of the present technology enable transverse position resolution, or "hit" resolution, of <NUM> or better in the disclosed tracking detectors. The continuous multiple scattering of a proton in matter limits spatial resolution. Measuring the proton transverse position before and after the patient can enable a typical uncertainty on the path through the patient as a function of depth of <NUM>, setting criterion for the hit resolution. Multiple scattering is more probable for lower energy protons, so a strategy of using lower energy protons to limit the residual range, while optimal for range resolution and practicality, has a potential drawback on spatial resolution. While range resolution is typically priority, a work-around for patients needing extra spatial detail would be to use a higher energy beam with additional passive material in front of the range detector. Resolution and dose trade-off would then be similar as for standard approaches.

In some examples, the residual range detector of the present disclosure comprises a rectangular volume of scintillator, with an array of large-area photomultiplier tubes mounted on the side downstream of the proton beam. The protons stop in the scintillator, generating a large number of scintillation photons within a few nanoseconds. To obtain fast signals, the sides not occupied by photomultiplier tubes are painted or covered with an anti-reflective material to absorb photons, and the photomultiplier tubes collect only direct photons that have not scattered off the walls. It is contemplated that signals from the photomultiplier tubes can be summed to produce a total energy signal, as well as weighted to produce an X-position signal and a Y-position signal. And, by recording only three signals per event (total energy, X-position, Y-position), a major advantage in electronics cost and data volume is achieved or realized. The high speed of the direct photon collection is another major advantage, compared to conventional designs with reflective surfaces that take much more time to collect the scintillation light and ultimately introduce errors in map reconstructions due to non-linearity in pulse height versus residual range trends. The monolithic design, combined with a strategy of limiting the residual range of the measurement, has the major advantages of reducing weight and optimizing dose. As an extra benefit, a position measurement is obtained from the residual range detector, adding extra redundancy useful for rejecting events with problems such as inelastic scatters.

Results have been obtained, as illustrated in at least <FIG>, with a residual range detector consisting of a 10x10x10 cm<NUM> active volume of plastic scintillator and a 2x2 array of photomultiplier tubes, and a pencil beam. The dependence of pulse height of the total energy signal on residual range is a consequence of the convolution of photon production and collection efficiency, and is approximately linear. The narrow spread in pulse heights at each range setting indicates that range resolution goals have been achieved. Additionally, as illustrated in at least <FIG>, a pulsed light emitting diode may be used for photomultiplier tube gain calibration. Such a technique may be used to maintain constant gains by comparing to a very stable photodiode.

The tracking system of the present disclosure, as illustrated in at least <FIG>, is based on scintillating fibers and multi-anode photomultiplier tubes. In some embodiments, <NUM><NUM> fibers are used, with two layers each for X and Y coordinates. Adjacent fibers in different layers are bundled into single photomultiplier channels, and protons may be detected in as few as one channel per view. Along with some backing material, the total width of each tracking plane or detector may be about <NUM>, and may provide a typical hit resolution of about <NUM>. A field size of <NUM>. <NUM><NUM> may allow entire field coverage with no gaps. Other examples are possible, and each example may be implementation-specific.

Using pencil beams and a position-sensitive range detector as disclosed allows for a reduction by a factor of twelve (<NUM>) the number of light sensors and electronics channels relative to conventional designs, by segmenting the tracker laterally, in orthogonal directions for X and Y, into strips <NUM> in width for example. Fibers are bundled, from the same position within different strips, into single photomultiplier tube anodes. The tracking system will precisely measure position within a strip, and information from the pencil beam settings or the proton radiograph will then indicate which strip the proton was in. In some embodiments, the total number of channels for a single X-Y tracking plane may be sixty-four (<NUM>) channels, and can be read out with a single multi-anode photomultiplier tube. Only the channels that are hit for a given event need to be recorded, based on a threshold detection algorithm. Pulse width may be approximately <NUM> nanoseconds. In comparison, for a <NUM> pencil beam, the average time between hits for a fiber in the beam will be about <NUM> nanoseconds. Therefore, a low overlap probability is achieved.

Thickness of tracking planes in terms of WEPL is a tradeoff, a worthwhile trade-off for the non-complex, fast system of the present disclosure which is scalable to large field sizes with no gaps between scintillating fibers. The trade-off involves a slightly higher dose for a given resolution: the relative dose increase is approximately the same as the relative material increase including the object to be imaged. For example, an extra <NUM> of tracking material will increase the dose to image a <NUM> object by <NUM>%. It may be possible to gain <NUM>% or <NUM>% with a more optimized design in the future. However, much larger gains are possible by optimizing the strategy for the residual range measurement as described above.

Some embodiments of the detector of the present technology are leveraged at average proton rates up to <NUM>, and measurements of ranges and positions of protons separated in time by as little as <NUM> nanoseconds or less. It is contemplated that if the detector cannot resolve events in adjacent bunches, these events will have to be discarded, slightly increasing the dose for a given image quality. Additionally, it is contemplated that the gain of the photomultiplier tubes should be high enough to obtain good signal-to-noise for low-range events, but low enough to remain within the current limits of the photomultiplier tubes for high-range (large pulse-height) events, especially when operating at <NUM>.

Some embodiments of the detector of the present technology are leveraged to demonstrate resolution per proton in WEPL of <NUM> or better across the sensitive detector area. This is a key specification to optimize proton range resolution relative to dose to the patient. A worse resolution will still be functional but will result in increased dose to the patient. An image will average many proton measurements to obtain a resolution of <NUM> or better. Stable photomultiplier gains may be important to achieve such resolution. Frequent and efficient calibration strategies, and alternative photomultiplier tube choices, may be considered.

Embodiments of the detector of the present technology may be leveraged to demonstrate proton detection efficiency of greater than <NUM>% per tracking plane, with transverse position resolution of <NUM> in order to maintain good spatial resolution. Undetected protons increase the dose without improving the image resolution. If the detection efficiency is not high enough, light yield may be increased through design. Alternatives include: using thicker scintillating fibers, with the drawback of adding additional material to the tracking detector(s); silicon photomultiplier sensors with large area are accessible, and may be an alternative to multi-anode photomultiplier tubes, with higher quantum efficiency.

Regarding calibration, LED pulsing maintains photomultiplier tube gains relative to a photodiode. Since photodiodes are very stable, with gain of <NUM>, the concept in <FIG> is very effective. Proton beam data is then used to calibrate the total energy signal, the x-position signal, and the y-position signal. Calibration data may be acquired in a three-dimensional grid of transverse position across the field and proton residual range. The data can be binned in a dimensional grid with coordinates of measured total energy signal, x-position signal, and y-position signal. For each bin, the average true energy, x-position, and y-position are stored. For any event, the dimensional grid can be used as a look-up table, interpolating if helpful, the true quantities from the measured quantities.

Regarding performance of the residual range detector of the present disclosure, range stacks intrinsically add to range straggling in a measurement since the measurement is of the stopping point of the proton. Similarly, a segmented calorimeter adds to range straggling from the material in segments before the segment in which the proton stops. The monolithic design of the residual range detector of the present disclosure, combined with a strategy of limiting the residual range of the measurement, has a number of advantages such as the possibility of the residual range measurement being limited by the range straggling when the proton exits the patient, rather than by the material of the range detector itself. Since delivered dose decreases as the square as the range resolution improves, this may have a significant impact. Also, limiting the residual range reduces the fraction of protons lost to nuclear scatters, again improving the dose performance. And, as an extra benefit, a position measurement in the residual range detector is obtained, adding extra redundancy useful for rejecting nuclear scatters for example.

Regarding proton radiography versus proton computerized tomography, proton radiography checks the range through the patient, while proton computerized tomography potentially measures directly the range to the tumor. Proton computerized tomography uses much more data, beam time, analysis, operational complexity, and a patient thin enough for imaging from all directions. Through conversations with clinicians, it is found that a practical range check would be very useful and is a higher priority than longer-term proton computerized tomography development. A strategy for proton radiography may include for example: continue to rely on an x-ray computerized tomography scan for treatment planning; for each planned field, prepare in advance simulations showing the expected proton radiograph from that direction compare to the actual proton radiograph for alignment and range check; if desired, range checks from additional directions in addition to the treatment direction can be done; if the patient is too thick in the treatment direction, a range check and alignment may be possible from another direction; if the patient is too thick from all directions, an alignment check is still possible using the edges of the patient; if the range check passes, an error in the range to the tumor would have to involve an unexpected cancellation, with extra material in front of the tumor matched by a deficit behind the tumor.

Various configurations may omit, substitute, or add various method steps or procedures, or system components as appropriate.

This description provides example configurations only, and does not limit the scope, applicability, or configurations of the claims. Rather, the preceding description of the configurations will provide those skilled in the art with an enabling description for implementing described techniques. Various changes may be made in the function and arrangement of elements without departing from the scope of the disclosure.

Also, configurations may be described as a process which is depicted as a flow diagram or block diagram. Although each may describe the operations as a sequential process, many of the operations can be performed in parallel or concurrently. In addition, the order of the operations may be rearranged. A process may have additional steps not included in the figure. Furthermore, examples of the methods may be implemented by hardware, software, firmware, middleware, microcode, hardware description languages, or any combination thereof. When implemented in software, firmware, middleware, or microcode, the program code or code segments to perform the tasks may be stored in a non-transitory computer-readable medium such as a non-transitory storage medium. In some examples, one or more processors perform the described tasks.

Furthermore, the example embodiments described herein may be implemented as logical operations in a computing device in a networked computing system environment. The logical operations may be implemented as: (i) a sequence of computer implemented instructions, steps, or program modules running on a computing device; and (ii) interconnected logic or hardware modules running within a computing device.

Claim 1:
A method for operating a medical imaging system (<NUM>), the method comprising:
generating a beam of particles;
steering the beam of particles through a first tracking detector (<NUM>), an object (<NUM>), a second tracking detector (<NUM>) and into a residual range detector (<NUM>; <NUM>), wherein the residual range detector includes at least one photon detector (<NUM>, <NUM>), wherein the first tracking detector includes first scintillating fibers, the first scintillating fibers are divided into a plurality of strips, wherein terminal ends from each first scintillating fiber at a first position in each of the plurality of strips are bundled together;
collecting tracking data from a first light detector (<NUM>), wherein the first light detector is coupled to the terminal ends of the bundled first scintillating fibers of the first tracking detector (<NUM>),
collecting tracking data from a second light detector (<NUM>), wherein the second light detector is coupled to the second tracking detector (<NUM>),
wherein the tracking data from the first light detector (<NUM>) and the tracking data from the second light detector (<NUM>) represent a trajectory of the beam of particles through the first tracking detector (<NUM>) and through the second tracking detector (<NUM>);
collecting energy data from the at least one photon detector, wherein the energy data represents energy loss of the beam of particles traversed through the object; and
generating an image of the object based on the tracking data and the energy data, characterized by during the steering of the beam of particles, varying an initial energy and transverse positions of the beam of particles through a range of values and at predetermined increments.