Patent Publication Number: US-10775460-B2

Title: Image guided radiation therapy system

Description:
The present invention relates generally to magnetic resonance guided radiotherapy treatment systems and, specifically, to an apparatus and method for magnetic resonance guided radiotherapy treatment with beam&#39;s eye view imaging. 
     BACKGROUND 
     Image guided radiotherapy (IGRT) has become the state of the art in radiation treatment. IGRT may utilize two-dimensional (2D) projection images or three-dimensional (3D) cone beam computed tomography (CT) images that are acquired prior to treatment. These images are compared to a set of pre-treatment images to ensure the patient has been set up accurately and consistently each treatment session. While CT, for example reconstructs in a standard cartesian geometry, the geometry of the radiation treatment beam originates from a divergent source. Beam&#39;s eye view (BEV) projection is commonly used to visualize patient anatomy to determine exactly what tissue will be irradiated from the divergent treatment beam. 
     In BEV projection, the 3D image dataset captured by CT, referred to the portal image, is registered to a reconstructed BEV image. The BEV image is reconstructed by ray tracing through the 3D CT dataset of the portal image from a virtual source position aligned with the location of the target of the radiation treatment beam source. Since the BEV image represents the path of the divergent radiation treatment beam, target coverage and critical structure avoidance can be accurately determined. This makes the BEV image an ideal image to use for real-time IGRT where the image must be acquired and analyzed in real time to reposition the beam to conform to the target volume while avoiding any surrounding radiation sensitive organs. However, processing the CT dataset in real time to generate a BEV image is computationally intensive and cannot be completed in real-time. 
     Magnetic resonance (MR) guided radiotherapy treatment systems integrate magnetic resonance imaging (MRI) devices with radiotherapy treatment systems. For example, U.S. Pat. No. 8,983,573, incorporated by reference in its entirety herein, is directed to a radiation therapy system that comprises a combined MRI apparatus and a linear accelerator capable of generating a beam of radiation. 
     In MRI, a signal from a slab of selected tissue gives rises to a two-dimensional slice that is integrated in a direction perpendicular to the slab. If the slab is oriented horizontally, each pixel in the 2D image is generated by summing the signal together along a vertical line. A radiation beam originating from a radiation source (a point source location) diverges from the source and fans out over the target. Thus, targeting the radiation source based on this conventionally obtained image may lead to a decreased radiation dose at the target than what was planned and/or unnecessary dose to tissue surrounding the target. For example, a 5 cm thick image slice may result in up to 4 mm of targeting error, 15 cm from the central beam axis, assuming a 100 cm distance from the radiation source. This effect will increase further as slice thickness increases. 
     While it may be possible to process the image data collected in MRI by ray tracing to produce an BEV image, similar to the methodology described above for CT, this processing is performed after the data is captured and therefore not available in real-time. Furthermore, the 3D image data collected in MRI requires significantly more data acquisition in general than does the image data collected in CT. As a result, significant time is required to produce a BEV image. 
     Accordingly, it is an object of the present invention to obviate or mitigate at least one of the above-noted disadvantages. 
     BRIEF SUMMARY 
     In accordance with an aspect of an embodiment, there is provided a magnetic resonance (MR)-radiotherapy (RT) hybrid system comprising: a radiation source configured to supply a radiation beam to treat the patient; and an MR imaging (MRI) apparatus configured to generate a divergent gradient field shaped to match a divergent geometry of the radiation beam of the radiation source. 
    
    
     
       BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING(S) 
       Embodiments of the invention will now be described by way of example only with reference to the following drawings in which: 
         FIG. 1  is an isometric view of an integrated radiation source and magnetic resonance imaging (MRI) system; 
         FIG. 2  is a view in a transverse plane of the integrated linear accelerator and MRI system of  FIG. 1 ; 
         FIG. 3  is a view in a saggital plane of the integrated linear accelerator and MRI system of  FIG. 1 ; 
         FIG. 4  is plan view of a transverse gradient coil for the MRI system of  FIG. 1 ; 
         FIG. 5  is an isometric view of another embodiment of a set of gradient coils for an MRI system using a cylindrical magnet; 
         FIG. 6  is a schematic view of a radiation beam used to target patient tissue based on images encoded from a magnetic field produced by a conventional gradient coil configuration; 
         FIG. 7  is a schematic view of the magnetic field produced by a conventional gradient coil configuration; 
         FIG. 8  is a schematic view of the radiation beam used to target patient tissue based on images encoded from a divergent magnetic field produced by a gradient coil configuration in accordance with an embodiment of the invention; 
         FIG. 9  is a schematic view of the divergent magnetic field produced by the gradient coil configuration in accordance with an embodiment of the invention; 
         FIG. 10  is an isometric view of another embodiment of an integrated radiation source accelerator and magnetic resonance imaging (MRI) system; 
         FIG. 11  is an isometric view of another embodiment of an integrated radiation source accelerator and magnetic resonance imaging (MRI) system; and 
         FIG. 12  is an isometric view of another embodiment of an integrated radiation source accelerator and magnetic resonance imaging (MRI) system. 
     
    
    
     DETAILED DESCRIPTION OF THE EMBODIMENTS 
     For convenience, like numerals in the description refer to like structures in the drawings. Referring to  FIGS. 1 to 3 , an integrated radiation source and MRI system is shown and is generally identified by reference numeral  110 . As can be seen, the integrated radiation source and MRI system  110  includes a radiation source and an MRI apparatus  114 . In this embodiment, the radiation source is a linear accelerator or linac  112 . As will be described, the linac  112  is configured to generate a treatment beam. The MRI apparatus  114  is configured to image a patient in real-time. The linac  112  and the MRI apparatus  114  are coupled to a rotatable gantry  122  so that they can be rotated in unison to treat a patient P. 
     In this particular example, the MRI apparatus  114  comprises a biplanar magnet having a pair of opposing magnet poles  118  and  120  creating a 0.2 T magnetic field strength. The MRI apparatus  114  is an open bore type including a table  116  on which the patient P can lay. In  FIG. 1 , the magnet poles  118  and  120  of the biplanar magnet are disposed above and below the table  116 . The linac  112  and magnet poles  118  and  120  are mounted on the gantry  122  that is supported by a frame  124 . In  FIG. 2 , the gantry  122  is rotated, and the magnet poles  118  and  120  of the biplanar magnet are disposed on the left-hand side and right-hand side of the table  116 , respectively. 
     The linac  112  includes a head  128  housing an electron beam generator  130  mounted on an arm  132  that is affixed to the gantry  122 . In this manner, the linac  112  rotates in unison with the gantry  122  and thus, maintains its position relative to the magnet poles  118  and  120 . If desired, the linac  112  may have its own gantry. In this case, the gantry of the linac  112  and the gantry  122  are mechanically coupled so that the linac  112  rotates in unison with the magnet poles  118  and  120 . 
     The electron beam generator  130  includes an electron gun  133 , an RF generator  134 , an accelerating waveguide  136 , a heavy metal target  138  at one end of the accelerating waveguide  136  and a beam collimating device (not shown). Interference reducing structure is also provided to inhibit the linac  112  and MRI apparatus  114  from interfering with one another during operation. 
     Alternatively, the linac  112  and MRI apparatus  114  may be mechanically coupled so that the electron beam is directed horizontally, and the magnet poles  118  and  120  are mounted vertically such that the magnetic field is horizontal, but perpendicular to the electron beam. These two components are fixed and non-movable. Variable angle electron or photon beam delivery is allowed by rotating the subject while in a sitting position. This integrated linac and MRI system configuration is particularly useful for lung cancer subjects who prefer standing/seating to laying supine, and for whom, conventional CT simulation does not allow simulation in the sitting position. 
     Further specifics of the integrated radiation source and MRI system  110  are described in Applicants&#39; U.S. Pat. Nos. 8,958,528 and 8,983,573, the contents of which are incorporated by reference. Additionally, other systems designed to integrate a linac with an MRI system may also be employed. 
     The MRI apparatus  114  comprises an MRI system that comprises a main magnet, gradient coils, a radio frequency (RF) coil and a scanner. The main magnet is configured to produce a magnetic field that aligns the hydrogen atoms of the patient P placed on the table with the direction of the magnetic field produced by the main magnet. In this embodiment, the main magnet is a biplanar magnet. The gradient coils are configured to produce a magnetic gradient field distribution that is generally weaker than the magnetic field produced by the main magnet. The magnetic gradient field distribution is superimposed on top of the magnetic field produced by the main magnet. The RF coil is configured to apply a RF pulse that is directed toward the area to be scanned. The RF pulse is configured to excite atoms to produce signals, which are then detected by RF receivers. The gradient coils are configured to localize the signal, first by defining the region of excitation affected by the RF pulse, and then by localizing the signal within a slice or slab of the area being scanned to generate images of the slide or slab. 
     The gradient coils are configured to produce linearly varying magnetic fields for spatial localization of the magnetic signal and image production. An example of winding patterns of gradient coils for the biplanar MRI system  114  is depicted in  FIG. 4 . In this example, the gradient coil varies the magnetic field in a direction perpendicular to the magnetic field produced by the main magnetic. The windings on the right and left sides of  FIG. 4  depict current loops that circulate in different directions (clockwise versus counter-clockwise). Gradient coils having the winding patterns illustrated in  FIG. 4  would be positioned with the current paths adjacent to pole plates above and below the patient P in the system  110 . 
     Although the MRI apparatus is described using biplanar magnets, cylindrical magnets may also be used. Gradient coil winding patterns for cylindrical MRI systems are illustrated in  FIG. 5 . As illustrated in  FIG. 5 , three sets of gradient coils may be employed. In such an embodiment, each gradient coil is driven by an independent power amplifier and creates a magnetic gradient field whose z-component varies linearly along the x-, y- and z-directions. The x and y gradient coils are transverse coils. The z gradient coil is a longitudinal gradient coil. The x and y gradient coils have a saddle (Golay) coil configuration. The z gradient coil has a circular (Maxwell) coil configuration. 
     As previously stated, the linac  112  includes a beam collimating device (not shown). In this embodiment, the beam collimating device is a multileaf collimator (MLC) that is configured to shape the treatment beam radiating from the linac  112 . 
     As described above, in conventional systems, the linac  112  is configured to aim the treatment beam based on MRI imaging using the conventional coordinate systems generated by the gradient coils of the MRI apparatus  114 . Referring to  FIG. 6 , the problem of making targeting decisions for the linac  112  based on MRI images created using conventional coordinate systems is illustrated. In the conventional coordinate system, targeting decisions are made on pixels that summed in a direction perpendicular to an image plane  702 . The images slices  704  used by the MRI apparatus to create an MRI image are in a direction perpendicular to a source-to-isocenter axis. However, the treatment beam  706  diverges as it travels from the linac. As a result of this divergence, a part  708  of the target tumor T may not be treated even though it will appear on the MRI image to be within a path of the treatment beam. Similarly, healthy tissue  710  may be unnecessarily radiated by the treatment beam even though it will not appear on the MRI image to be within the path of the treatment beam. As will be appreciated, such divergence results in incomplete treatment of the patient&#39;s condition as well as significant radiation of healthy tissue which can have adverse side effects. 
     The in-plane encoding gradient of the MRI apparatus  114  is given by Equation 1.
 
 G   i ( x,y,z )∝ {circumflex over (r)}   i   ·     x,y,z         Equation 1
 
     Where {circumflex over (r)} i  represents a unit vector identifying one of the encoding directions within the plane of the imaging slice.  FIG. 7  depicts a qualitative pattern of the encoding field distribution of the magnetic field B z  produced by gradient coils, which are used to allow for spatial encoding within the plane of the imaging slab. As will be appreciated, the field variation is the same at all “vertical” positions within the imaging slab. 
     In accordance with an embodiment, the gradient coils of the MRI apparatus  114  are configured to generate a divergent magnetic gradient field shaped to match the divergent geometry of the treatment beam emanating from the radiation source. Thus, the radiation beam is considered to be emanating from a point source. As such, the divergent geometry of the treatment beam will depend on the distance of the radiation source from the imaging isocenter. The divergent magnetic gradient field is shaped such that image pixels scanned by the RF detector are summed over the same divergent path as the treatment beam produced by the linac  112 . Targeting decisions for the linac  112  are made based on pixels acquired from the gradient coils of the MRI apparatus  114  which images the patient using a divergent perspective as per the divergent magnetic gradient field. 
       FIG. 8  illustrates benefits of making targeting decisions for the linac  112  based on MRI images created using the divergent coordinate system of the gradient coils of the MRI apparatus  114 . In the divergent coordinate system, targeting decisions are made on image pixels that summed in a direction in-line with the divergent direction of the treatment beam of the linac  112 . That is, the image slices used by the MRI apparatus  114  to create an MRI image mimic the shape of the treatment beam. As a result, the treatment beam is more likely to radiate all of the targeted tissue of the tumor, while minimizing radiation of healthy tissue adjacent to the tumor. Thus, the using of a divergent gradient magnetic field minimizes adverse side effects and more completely treats the target tumor. 
     The in-plane encoding gradient of the MRI apparatus  114  is given by Equation 2. 
     
       
         
           
             
               
                 
                   
                     
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     As previously stated, {circumflex over (r)} i  represents a unit vector identifying one of the encoding directions within the plane of the imaging slice. In a BEV image, this will typically be one of the in-plane encoding directions. In equation 2, {circumflex over (r)} s  represents a unit vector identifying a direction of the axis connecting the radiation source and imaging isocentre (referred to as the source-to-isocenter axis). In a BEV image, this will typically be the vector perpendicular to the image plane. Furthermore, SID represents the distance between the radiation source and the magnetic isocenter of the MRI apparatus  114 . 
     In this embodiment, the in-plane encoding gradient is altered from the encoding gradient of Equation 1 to equal or approximate the encoding gradient of Equation 2 by altering the winding patterns of the gradient coils of the MRI apparatus  114  to produce a divergent coordinate system. 
     The gradient coils are configured to generate a divergent magnetic gradient field shaped to match the radiation beam of the radiation source. In one embodiment, the gradient coils are produced and designed in a similar manner to conventional gradient coils, but are designed with altered winding patterns to generate the required divergent gradient field distributions in line with the treatment beam. These winding patterns are typically derived numerically through a variety of established methods and can be optimized to approximate any desired field pattern. For example, Poole M., Bowtell R. Novel gradient coils designed using a boundary element method. Concepts in Magnetic Resonance Part B: Magnetic Resonance Engineering. 2007 Aug. 1; 31(3):162-75, which is incorporated herein by references, describes methods of deriving winding patterns. However, such a configuration would permanently acquire images only in a divergent frame. 
     If it is desired to provide the system with extra flexibility, additional gradient coils can be added to the conventional gradient coils. In an aggregate configuration, the additional gradient coils are configured to modify the transverse gradient field to the desired divergent magnetic field. Alternatively, in a separate configuration, the additional gradient coils are configured to provide the desired divergent magnetic field on their own. In both the aggregate and separate configurations, the additional gradient coils may be controlled independently from the conventional gradient coils. 
     Thus, if the additional gradient coils are not activated, MRI imaging is performed using the conventional gradient coils in a standard, transverse Cartesian coordinate system. In the aggregate configuration, if the additional coils are activated, MRI imaging is performed using a combination of the additional coils and the conventional coils to provide the divergent coordinate system. In the separate configuration, if the additional coils are activated, MRI imaging is performed using the additional coils alone in order to provide the divergent coordinate system. 
     In a system in which the gradient coils and radiation source are configured to rotate in unison, such as the one described above, the minimum number of additional coils to generate a divergent geometry in both in-plane encoding directions is two; one coil for each of the two axes perpendicular to the source-to-isocenter axis. As will be appreciated by one of skill in the art, additional coils could be used. 
     A qualitative pattern of the encoding field distribution B z  of the gradient coils of the MRI apparatus  114  is displayed in  FIG. 9 . The magnetic field B z  produced by gradient coils of the MRI apparatus  114  is used to allow for spatial encoding within the plane of the imaging slab. The in-plane encoding gradient is configured for Beam&#39;s Eye View (aligned with the linac  112  treatment geometry) imaging. The field pattern has a stronger variation at some “vertical” positions within the slab than at others, allowing the imaging to inherently represent a divergent geometry. 
     In the embodiments described above, the gradient coils and the linac are fixed to a gantry so that they rotate in unison. However, a system in which the gradient coils do not rotate with the radiation source will also benefit from the generating a divergent gradient coordinate system described above. However, it may not be possible to design gradient coils that generate divergent magnetic gradient fields shaped to match all possible positions of the treatment beam, as the direction of the treatment beam is constantly changing with the rotation of the radiation source. However, the desired magnetic gradient fields can be effectively approximated by an optimized combination of a set of additional coils that generate non-linear magnetic field distributions in addition to the conventional gradient coils. 
     The additional set of coils can be the “2 nd  order” field gradients that are often used for shimming. These coils consist of a winding pattern that produces B z  field patterns that correspond to the second-order spherical harmonic functions. While many magnets already have such sets of coils, they are typically not engineered to handle the power requirements of imaging, and would have to be specifically designed for this function by generating winding pathways with different criteria such as lower overall resistance and inductance. 
     There are five of spherical harmonic functions, and a coil is required to produce each of them. If a divergent geometry is requited for any arbitrary image orientation in the magnet geometry, five coils may be necessary to approximate the field relationship in Equation 2 in conjunction with the conventional gradient coils. However, only four coils would be required if it can be assumed that the BEV image plane will always be parallel to the axis of gantry rotation. This assumption regarding the orientation of the BEV image plane would allow one of the five harmonic fields to be omitted. Thus, only four coils are required. 
     While a particular integrated radiation source and MRI system has been described, one of skill in the art will appreciate that other configurations are possible. Turning now to  FIG. 10 , another embodiment of an integrated radiation source and MRI system is shown and is generally identified by reference numeral  210 . As can be seen, the integrated radiation source and MRI system  210  comprises a radiotherapy radiation source  213 , a biplanar MRI magnet assembly  214 , a patient treatment couch  216  and a rotating gantry  222 . The radiotherapy radiation source  213  is configured to generate a treatment beam. The MRI magnet assembly  214  is configured to image a patient positioned on the patient treatment couch  216  in real time. The radiotherapy radiation source  213  and the MRI magnet assembly  214  are coupled to the rotating gantry  222  which is supported by a frame such that the radiation source  213  and the magnet assembly  214  can be rotated in unison to image and treat the patient. 
     In this particular embodiment, the radiation source  213  comprises a collimating device  224  that is configured to shape the treatment beam radiating from the radiation source  213  to treatment the patient on the treatment couch  216 . The axis defined by the treatment beam generated by the radiation source  213  is generally parallel to the axis defined by the poles  218  and  220  of the magnet assembly  214 . 
     In this particular embodiment, the biplanar MRI magnet assembly  214  comprises a biplanar magnet having a pair of opposing biplanar magnet poles  218  and  220 , and a biplanar gradient coil set  223 . At least pole  218 , the pole closest to the radiation source  213 , includes a centrally located open bore sized to allow the treatment beam generated by the radiation source  213  to pass through and treat a patient on the patient treatment couch  216 . While the poles  218  and  220  are shown in  FIG. 10  as being above and below the patient, as the gantry  222  rotates around the patient treatment couch  216  the biplanar MRI magnet  214 , the poles  218  and  220  may move 360 degrees around the treatment couch  216 . The gradient coil set  223  of the magnet assembly  214  is configured to generate a divergent magnetic gradient field shaped to match the divergent geometry of the treatment beam of the radiation source  213 . The treatment beam generated by the radiotherapy radiation source  213  is configured to pass through the open bore of the pole  218  of the MRI magnet assembly  214  during operation. 
     Turning now to  FIG. 11 , another embodiment of an integrated radiation source and MRI system is shown and is generally identified by reference numeral  310 . As can be seen, the integrated radiation source and MRI system  310  comprises a radiotherapy radiation source  313 , a biplanar MRI magnet assembly  314 , a patient treatment couch  316  and a rotating gantry  322 . The radiotherapy radiation source  313  is configured to generate a treatment beam. The MRI magnet assembly  314  is configured to image a patient positioned on the patient treatment couch  316  in real time. The radiotherapy radiation source  313  and the MRI magnet assembly  314  are coupled to the rotating gantry  322  which is supported by a frame such that the radiation source  313  and the magnet assembly  314  can be rotated in unison to image and treat the patient. 
     In this particular embodiment, the radiation source  313  comprises a collimating device  324  that is configured to shape the treatment beam radiating from the radiation source  313  to treatment the patient on the treatment couch  316 . The axis defined by the treatment beam generated by the radiation source  313  is generally perpendicular to the axis defined by the poles  318  and  320  of the magnet assembly  314 . 
     In this particular embodiment, the biplanar MRI magnet assembly  314  comprises a biplanar magnet having a pair of opposing biplanar magnet poles  318  and  320 , and a biplanar gradient coil set  323 . While the poles  318  and  320  are shown in  FIG. 11  as being above and below the patient, as the gantry  322  rotates around the patient treatment couch  316  the biplanar MRI magnet  314 , the poles  318  and  320  may move in 360 degrees around the treatment couch  316 . The gradient coil set  323  of the magnet assembly  314  is configured to generate a divergent magnetic gradient field shaped to match the divergent geometry of the treatment beam of the radiation source  313 . 
     Turning now to  FIG. 12 , another embodiment of an integrated radiation source and MRI system is shown and is generally identified by reference numeral  410 . As can be seen, the integrated radiation source and MRI system  410  comprises a radiotherapy radiation source  413 , a cylindrical MRI magnet assembly  415 , a patient treatment couch  416  and a rotating gantry  422 . The radiotherapy radiation source  413  is configured to generate a treatment beam. The MRI magnet assembly  415  is configured to image a patient positioned on the patient treatment couch  416  in real time. 
     In this particular embodiment, the radiation source  413  comprises a collimating device  424  that is configured to shape the treatment beam radiating from the radiation source  413  to treatment the patient on the treatment couch  416 . The treatment beam generated by the radiation source  413  passes between elements of the cylindrical magnet assembly  416  as shown in  FIG. 12 . The radiation source  413  is coupled to the gantry  422  such that the radiation source  413  can treat any party of the patient on the patient treatment couch  416  through rotation of the gantry  422 . In this particular embodiment, the radiation source  413  comprises a collimating device  424  that is configured to shape the treatment beam radiating from the radiation source  413  to treatment the patient on the treatment couch  416 . 
     In this particular embodiment, the cylindrical MRI magnet assembly  414  comprises a pair of cylindrical magnets and a cylindrical gradient coil set  425 . The gradient coil set  425  comprises two gradient magnets which are within the cylindrical magnets. The axis defined by the treatment beam generated by the radiation source  413  is generally perpendicular to the axis defined by the cylindrical magnets of the cylindrical MRI magnet assembly  414 . The gradient coil set  425  of the magnet assembly  414  is configured to generate a divergent magnetic gradient field shaped to match the divergent geometry of the treatment beam of the radiation source  413 . 
     Although embodiments of the invention have been described herein, it will be understood by those skilled in the art that variations may be made thereto without departing from the scope of the appended claims.