A radiotherapy device

Abstract

Disclosed herein is a linear accelerator for a radiotherapy apparatus, the linear accelerator comprising: an acceleration waveguide 104 having a geometrical structure arranged to support at least two

Description

This disclosure relates to an acceleration waveguide for a linear accelerator and in particular to an acceleration waveguide for a linear accelerator for a radiotherapy device. Background Radiotherapy can be described as the use of ionising radiation to damage or destroy unhealthy cells in both humans and animals. The ionising radiation may be directed to tumours on the surface of the skin or deep inside the body. Common forms of ionising radiation include x-rays and charged particles. A radiotherapy device typically comprises a gantry which supports a beam generation system, or other source of radiation, which is rotatable around a patient. Alternatively, static, fixed beam systems may also be used. The beam generation system is typically based on a particle accelerator such as a linear accelerator or 'linac' which comprises a radio frequency (RF) power source, a charged particle source, and an RF cavity. Linear accelerators (especially those for medical use) accelerate charged particles such as electrons to relativistic speeds along an acceleration path through an acceleration waveguide. The acceleration waveguide comprises one or more resonant cells that surround the charged particle trajectory. The acceleration waveguide is filled with RF power from an RF power source which forms an oscillating electric field, or an electromagnetic (EM) wave, inside the cavity. Charged particles are injected from a particle source such as an electron gun into the cavity, forming a beam. As the beam traverses the oscillating EM wave, it gains energy and is thus accelerated, often up to relativistic speeds. Accelerated particles with increased effective mass deposit more energy when they collide with other matter, which is usually either biological tissue or a tungsten target, the target being used to generate x-rays. The resultant particle beam, or x-rays, may be used for imaging or treatment, so often the energy of the beam is variable. Currently, separate imaging systems attached to the radiotherapy apparatus are utilised for pre- and post- treatment imaging, as well as imaging 'on-line' during treatment. This presents numerous engineering difficulties, along with adding expense and complexity to the design and operation of the apparatus. Additionally, the imaging source and associated detector have to be rotated along with the treatment source, further complicating the radiotherapy process. It can be advantageous to have a kV energy imaging source that is co-incident with an MV energy treatment beam source. However, such a co-incident set up cannot be implemented with a separate imaging device. A possible approach is to accelerate electrons along a linear accelerator into one of two mechanically moveable targets, with each target producing a different x-ray energy. A mechanism is used to mechanically switch between the two targets depending on the energy of the beam required. However, this adds an additional step to the process of transitioning from a treatment beam to an imaging beam, which also further complicates the radiotherapy process. It would thus be advantageous if there were a linear accelerator able to produce each of the respective beam energies required for imaging and radiotherapy treatment without the need for additional complications to the radiotherapy process. The present disclosure seeks to address this, and other disadvantages encountered in the prior art. Summary An invention is set out in the independent claims. Optional features are set out in the dependent claims. Examples will now be described, by way of example, with reference to the drawings of which: Fig. 1 shows a radiotherapy device or apparatus; Fig. 2 shows a beam generation system for a radiotherapy device; Fig. 3 shows components of an acceleration waveguide for a linear accelerator; Fig. 4a shows properties of a treatment mode of an acceleration waveguide; Fig. 4b shows properties of an imaging mode of the acceleration waveguide; Fig. 5 depicts an example of an RF power input system; Fig. 6 is a flowchart diagram which shows a method of operation of a linear accelerator; Fig. 7 shows a block diagram of one implementation of a radiotherapy system; Fig. 8 shows a computer readable medium or, more generally, a computer program product. Overview The present disclosure relates to a linear accelerator designed such that it can produce electron beams of different energies, such that the linear accelerator can be utilised for both imaging and radiotherapy treatment purposes. Linear accelerators in radiotherapy produce therapeutic beams of x-ray radiation with energies in the MV range, whereas the x-ray beam energies required for imaging purposes are in the kV range. It would be desirable to use the same linear accelerator for both imaging and therapeutic purposes, as it would provide a co-incidental imaging source during radiotherapy treatment, without the complexity of requiring an additional step of adjusting the target. However, use of the same linear accelerator for imaging and therapeutic purposes has been inhibited by engineering / design features in the past. For example, adding energy switches to a linear accelerator (to enable the accelerator to change the energy of the accelerated electrons) introduces mechanical complications and complications related to maintaining the vacuum of the accelerator. In order for a linear accelerator to produce x-rays of a particular energy, the electron beam must have an energy which closely matches the energy of the desired x-rays. The energy of the electron beam is a direct consequence of the electromagnetic field in the acceleration waveguide of the linear accelerator, and the acceleration waveguide of the linear accelerator is typically carefully designed around an energy of the beam which is desired. In relation to a linear accelerator design, a waveguide mode can be thought of as the distribution of the electromagnetic field across the cross-section of the acceleration waveguide. Typically, the design of a linear accelerator is optimised for one fundamental mode, and thus one beam energy. However, both lower order modes (LOM) and higher order modes (HOM) can be excited in the acceleration waveguide in addition to the fundamental mode. These additional modes are generally seen as problematic, as they have unintended effects on the electron beam which are not conducive to beam acceleration. Much attention has been paid to eliminating these higher and lower order modes and typically these higher and lower order modes are avoided and / or eliminated using damping techniques. The present disclosure concerns a linear accelerator acceleration waveguide which is designed to specifically utilise a lower order and / or higher order mode to produce x-rays in the kV range, which are suitable for use in imaging. In some examples, the acceleration waveguide is designed such that one of a LOM or HOM or a combination of a LOM and / or a HOM may be used to generate a field which is conducive to kV imaging, whereas the fundamental mode of the linear accelerator can still be used to produce a treatment beam. This can be facilitated through the use of an RF source which can provide multiple frequencies, one of which relates to the fundamental mode, and one or more of which will result in at least one HOM and / or at least one LOM, giving the desired imaging mode(s). There exist a number of complexities when it comes to designing a linear accelerator such that it can utilise a HOM or LOM. During the process of optimising (i.e. tuning) the acceleration waveguide of the linear accelerator for a particular mode, the other possible modes (HOMs and LOMS) do not remain constant, rather they also are 'tuned' or adjusted in a non-linear fashion. Optimising a linear accelerator to operate at more than one mode becomes a non-linear system problem, which becomes quite complex to solve. Such a problem can be solved by using a Pareto front approach which includes an additional Pareto front for each additional HOM and / or LOM which is intended for use as an imaging mode. Such an approach can be used to simulate, design, identify and / or tune an acceleration waveguide having a geometrical structure arranged to support at least two modes of electron acceleration, such as a fundamental mode for treatment and a HOM and / or LOM for imaging. In addition, it is desirable to design the acceleration waveguide such that the frequency of the LOM or HOM can be driven with the same RF power source as the fundamental operating mode. This is a challenge, as typical RF power sources used to drive linear accelerators, e.g. a magnetron, can have a very limited bandwidth of frequencies to choose from. Thus, in order to utilise a fundamental mode for acceleration purposes, and a LOM or HOM (or combination thereof) for imaging purposes, the frequencies required for each purpose should preferably be within the bandwidth available from a suitable RF power source. Issues with electron bunching can occur when the electromagnetic field in the acceleration waveguide is not the electromagnetic field for which it was designed. Increasing the power delivered to the acceleration waveguide can increase the energy produced by the acceleration waveguide, which in turn increases the energy delivered to the electrons. However, this is typically ineffective at increasing the energy of the electron beam due to "bunching" complications. Typical acceleration waveguide designs involve one or more "buncher" sections which function to group the injected electrons into bunches for further acceleration in at least one subsequent "relativistic" section of the acceleration waveguide. Conventional "buncher" part(s) of the linear accelerator are designed around a single electron beam energy, to the extent that the "bunchers" may stop working altogether if the electrons are not at the required energy. This further complicates the design of a single linear accelerator which can be used to produce x-rays of varying energy suitable for both imaging and radiotherapy treatment. In the present disclosure, an acceleration waveguide and / or linear accelerator for a radiotherapy apparatus is provided that can utilise one, or a combination of, higher order or lower order modes in order to produce the distinct beam energies required for both imaging and radiotherapy treatment. Disclosed herein is a linear accelerator that comprises an acceleration waveguide having a geometrical structure arranged to support at least two modes of electron acceleration. The linear accelerator also comprises a target configured to generate x-rays for radiotherapy when impacted by electrons accelerated by a first of the at least two modes of electron acceleration, and to generate x-rays for imaging when impacted by electrons accelerated by a second of the at least two modes of electron acceleration. Advantageously, the approaches disclosed herein provide a linear accelerator capable of switching between treatment and imaging by using energy switching and / or energy selection. Furthermore, the approaches disclosed herein are capable of achieving energy switching and / or selection using components that may be fully contained within the conventional vacuum applied to a linear accelerator, and thus do not introduce a risk of interfering with that vacuum. Using the approaches disclosed herein, energy switching may be performed very quickly, because no mechanical switching components are required. Detailed Description Fig. 1 shows an exemplary radiotherapy (RT) device 100. The device and its constituent components will be well known to the skilled person but is described here generally for the purpose of providing useful accompanying information for the present disclosure. The radiotherapy device 100 is based on a linear accelerator (linac). The device shown in Fig. 1 combines magnetic resonance (MR) imaging capability with a linac-based radiotherapy capability, and is known as an MR-linac device. However, the present disclosure may be implemented in any radiotherapy device, for example, a linac-based radiotherapy device without magnetic resonance imaging capability. In operation, the MR scanner produces MR images of the patient, and the RT apparatus produces and shapes a beam of radiation and directs it toward a target region within a patient's body in accordance with a radiotherapy treatment plan. The MR-linac device 100 shown in Fig. 1 comprises an RF power source 102, an RF transmission apparatus 103, an acceleration waveguide 104, an electron source 106, a treatment head including a collimator 108 such as a multi-leaf collimator used to shape a treatment beam 110, MR imaging apparatus 112 (shown partially cut away), and a patient support surface 114. The RF transmission apparatus 103 comprises a waveguide component, which may be a copper waveguide. The depicted device does not have the usual 'housing' which would cover the MR imaging apparatus and RT apparatus in a commercial setting such as a hospital. In use, the device would also comprise the housing, part of which, together with the ring-shaped gantry, defines a bore. In particular, a part of the housing encloses the inner surface of the ring-shaped gantry, defining a bore through the device 100. The patient support surface 114 is moveable and can be used to support a patient and move them, or another subject, into the bore when an MR scan and / or when radiotherapy is to commence. The MR imaging apparatus 112 is configured to obtain images of a subject positioned on the patient support surface 114. The MR imaging apparatus 112 may be conventional MR imaging apparatus operating in a known manner to obtain MR data, for example MR images. The skilled person will appreciate that such MR imaging apparatus 112 may comprise a primary magnet, one or more gradient coils, one or more receive coils, and an RF pulse applicator. The RT device has a beam generation system comprising the RF power source 102, the acceleration waveguide 104, and the electron source 106. The beam generation system is configured to produce a beam of ionising radiation, otherwise known as the treatment beam 110, that is collimated and shaped by the collimator 108 and directed towards the bore. Typically, a radiation detector is positioned diametrically opposed to the collimator. The radiation detector is suitable for, and configured to, produce radiation intensity data. In particular, the radiation detector is positioned and configured to detect the intensity of radiation which has passed through the subject. The radiation detector may form part of a portal imaging system. The beam generation system is attached to the rotatable gantry 116 so as to rotate with the gantry 116. In this way, the beam generation system is rotatable around the patient so that the treatment beam 110 can be applied from different angles around the gantry 116. In a preferred implementation, the gantry is continuously rotatable. In other words, the gantry can be rotated by 360 degrees around the patient, and in fact may continue to be rotated past 360 degrees. The gantry is ring-shaped, i.e. a ring-gantry. The device 100 of Fig. 1 is controlled by a controller (not shown). The controller is a computer, processor, or other processing apparatus. The controller may be formed by several discrete processors; for example, the controller may comprise an MR imaging apparatus processor, which controls the MR imaging apparatus 112; an RT apparatus processor, which controls the operation of the RT apparatus; and a subject support surface processor which controls the operation and actuation of the patient support surface 114. The controller is communicatively coupled to a memory, e.g. a computer readable medium. Fig. 2 shows an example beam generation system 200 that will be described generally for the purpose of providing useful accompanying information for the present disclosure. For example, the beam generation system 200 may be used as the beam generation system in the device 100 of Fig. 1. The beam generation system 200 is based on a linear accelerator design, and may alternatively be referred to as a linear accelerator. The beam generation system 200 comprises an acceleration waveguide 202 and a source 204 of electrons. The source 204 of electrons may be an electron gun, for example a triode electron gun or diode electron gun. The acceleration waveguide 202 is configured to accelerate particles, in this case electrons, along an acceleration path 206 into a target 208, in order to produce a treatment beam 210 of radiation. The acceleration path 206 is also known as the central beam axis of the acceleration waveguide 202. The acceleration waveguide 202 comprises a series of cells. In this example, each cell has substantially the same shape and dimensions, but in other examples, that may not be so. The cells may be arranged such that each cell is RF-uncoupled / independent, and in that case each cell functions as a separate resonant cavity. In other implementations, such as the example of Fig. 2, the cells may be coupled together and, in that case, the overall coupled structure may be considered to be a single resonant cavity. In such an implementation, although the coupled cells function as a single resonant cavity, individual cells may still be referred to as cavities by those skilled in the art. The acceleration path is coincident with the centre axis of the acceleration waveguide 202 and passes through an aperture at the centre of each cell. The acceleration waveguide 202, the source 204 of electrons, the cells and the target 208 are enclosed within an evacuated and vacuum-sealed casing 212 to ensure that propagation of the electrons is not impeded as they travel toward the target 208. The vacuum-sealed casing 212 is evacuated using a vacuum system to ultra-high vacuum (UHV) conditions. As the electrons are accelerated in the acceleration waveguide 202, in some embodiments the electron beam path may be controlled by a suitable arrangement of steering magnets, or steering coils (not shown), which surround the acceleration waveguide 202. The arrangement of steering magnets may comprise, for example, two sets of quadrupole magnets. A source of RF waves, or RF power source 214, such as a magnetron or a klystron, is configured to produce and / or amplify RF waves. The RF power source 214 is coupled to the acceleration waveguide 202 via an RF transmission apparatus 216, which usually comprises copper waveguide sections that can have a circular or rectangular cross section. A modulator is configured to pulse RF waves through the copper waveguide into the acceleration waveguide 202. Typically, the RF waves are input into a particular cell of the acceleration waveguide 202. The RF transmission apparatus 216 that connects the RF power source 214 to the input cell of the acceleration waveguide 202 may comprise a waveguide network and may contain an RF window which may separate a vacuum envelope from an SF6 envelope. The RF transmission apparatus 216 is perpendicular to the acceleration waveguide 202 central beam axis 206 where it couples the power into the input cell. The RF input connecting pipe or tube is coupled with the acceleration waveguide 202 and joins the acceleration waveguide 202 at a substantially 90° angle. The RF transmission apparatus 216 may include a circulator 218 of any appropriate known type. The beam generation system can operate with either a standing wave or a traveling wave configuration. In a standing wave configuration as shown, the RF power source 214 is configured to pulse RF waves into the acceleration waveguide 202, in order to set up a standing wave of varying electric field that is suitable for accelerating charged particles. Although the RF power source 214 can operate in continuous mode, typically it operates in pulsed mode in view of the RF power levels required. An example RF wave frequency is 3 GHz, with a pulse duration in the range of microseconds and a pulse repetition rate in the range of several hundred pulses per second. The RF power source 214 may be a commercially available magnetron such as an E2V 3.1 MW magnetron, or any standard radiotherapy magnetron, operating at 3 GHz. Typically, the RF power source 214 produces each pulse at a particular phase in order to improve the stability of the standing wave within the acceleration waveguide 202. After it has been pulsed into the acceleration waveguide 202, some of the RF energy dissipates into the walls of the acceleration waveguide 202. In an acceleration waveguide made up of coupled cells, the standing RF wave is established according to the resonant frequency of the coupled structure. An effect of coupling individually resonant cells together to form a single resonant cavity is that, due to dispersion, a band of different frequency oscillation modes comprising higher and lower order modes may be permitted within the acceleration waveguide 202 either side of the resonant frequency of the coupled structure. The frequency of the RF waves provided by the RF power source 214 determines the mode(s) that are excited in the acceleration waveguide 202. There are also multiple modes of operation by which a standing wave at the resonant frequency can accelerate electrons within the acceleration waveguide. Electrons will accelerate or decelerate depending upon the polarity of the electric field they experience. The length of each cell in the cavity is designed such that the beam sees the same phase of the RF in each cell. The beam is synchronised such that on each oscillation the beam interacts with the positive part of the wave and is accelerated further. In one operational mode, known as the zero mode, the electric field of the standing wave has the same polarity and magnitude in all cells at any given time. During the time that an electron takes to traverse a given cell and enter the next cell, the field makes a complete oscillation, for example from positive to negative and back to positive, such that the electron sees the same accelerating field it has just experienced, rather than a decelerating field. Alternatively, a 'n mode' may be used. Rather than the electric field being of the same polarity in each cell at a given time, adjacent cells have opposite polarities at a given time. However, the dimensions of each cell are such that during the time an electron takes to traverse a given cell, the adjacent cell experiences a half oscillation in field polarity such that the electron entering the adjacent cell experiences an accelerating field polarity rather than a decelerating one, and so on. The source 204 of electrons, such as an electron gun, is also coupled to the acceleration waveguide 202 and is configured to inject electrons into the acceleration waveguide 202. The injection of electrons into the acceleration waveguide 202 is synchronised with the pulsing of the radiofrequency waves into the acceleration waveguide 202. In some implementations, an upstream portion of an acceleration waveguide in a linac may be referred to as a buncher section. The buncher section may comprise one or more cells of the acceleration waveguide. Within the buncher section, the phase of the RF wave, whether a standing wave or traveling wave, decelerates some electrons to allow slower electrons to catch up, concentrating the electrons in bunches. The electrons are then free to move together in so called "packets" or "bunches" and the bunches quickly accelerate to relativistic speeds through the subsequent cells of the acceleration waveguide. The acceleration waveguide may be designed with a buncher section that is optimised to produce an electron beam with a particular energy and intensity by bunching electrons into a beam of short pulses. RF waves may be input to the acceleration waveguide at a particular cell, or at more than one cell. In particular, RF waves may be input at a cell that is adjacent to the buncher portion of the acceleration waveguide. In the example of Fig. 2, the first two cells on the left-hand side of the acceleration waveguide 202 are the buncher and the following cells act to accelerate the electrons to relativistic speeds. Alternatively, the RF waves may be input into one or more of the cells belonging to the buncher section of the acceleration waveguide 202. Once the electrons have been accelerated to faster energies, such as 8 MeV or 10 MeV, they may pass into a flight tube. The flight tube is connected to the acceleration waveguide by a connecting tube. The flight tube is also kept under vacuum conditions. This connecting tube or connecting structure is termed a drift tube. The drift tube also forms part of a vacuum tube along with the other components within the vacuum-sealed casing 212. The electrons may travel along a slalom path toward the heavy metal target. Whilst the electrons travel through the flight tube, an arrangement of focusing magnets act to direct and focus the beam on the target. The slalom path allows the overall length of the linac to be reduced while ensuring that the beam of accelerated electrons, which is comprised of electrons with a small spread of energies, is focused on the target. The electrons travel toward the target 208 which may comprise, for example, tungsten, or another heavy metal. The impact of the electrons on the target 208 produces x-rays which form the treatment beam 210. When the electrons strike the target 208, x-rays are produced in a variety of directions. A primary collimator may block x-rays travelling in certain directions and pass only forward travelling x-rays to produce the treatment beam 210. The x-rays may be filtered and may pass through one or more ion chambers for dose measuring. The beam can be shaped in various ways by beam-shaping apparatus, for example by using the multi-leaf collimator 108, before it passes into the patient as part of radiotherapy treatment. If a flight tube is used, the target is located inside the flight tube and is located at the end of the flight tube to seal the vacuum system. The flight tube also comprises a target window, which is transparent to x-rays, and which is positioned to allow the x-rays which are produced when the beam generation system is in operation to pass from the evacuated flight tube through the target window and into the treatment head. In some implementations, the electrons are accelerated within an acceleration waveguide by using a travelling wave rather than a standing wave. In this case electrons travel at the phase velocity of the travelling wave, accelerated by the longitudinal electric field component. The acceleration waveguide 202 must be designed such that the phase velocity of the traveling wave does not exceed the speed of light, otherwise no acceleration of electrons will occur. In particular, using a disk-loaded waveguide, rather than a cylindrical waveguide, reduces the phase velocity appropriately such that electrons are accelerated. For an accelerator that uses a traveling wave, in addition to an RF input, the acceleration waveguide will have an RF output configured to transfer RF energy out of the acceleration waveguide and prevent it from reflecting and establishing a standing wave. If a drift tube is used adjacent to the acceleration waveguide, the RF output may be coupled to the drift tube. As with the input transmission apparatus or waveguide, which introduces RF power to the acceleration waveguide, the output waveguide through which RF power exits the waveguide can be connected via an elbow joint or 7-shaped' joint. RF waves pass out from the evacuated system via an RF output window which seals the vacuum envelope. Fig. 3 shows a cross section view along the longitudinal axis of an acceleration waveguide 300 suitable for use in a particle accelerator, for example for use as the acceleration waveguide 202 of the beam generation system 200 of Fig. 2, and depicts a typical multiple cell cavity. The acceleration waveguide 300 is suitable for use in a linac as shown in Fig. 1 and Fig. 2, but also could be used in other accelerators (e.g. a curved accelerator such as a synchrotron). The below examples and discussion relate to the acceleration of electrons, but the cavity can be used to accelerate any charged particle and therefore in any charged particle accelerator. For example, protons, positrons and ions can be accelerated using the techniques described herein. Two cells 310 of a series of connected cells are shown. The cells may be coupled together. The cells are each connected along a central axis 312 by irises 314, 315. Only two cells are shown in Fig. 3, although a typical acceleration waveguide will have more. The precise number will vary, dependent on the design criteria of the accelerator. Each cell is defined in the form of a recess within a surrounding shell of a conductive material, usually copper. The acceleration waveguide 300 may be described as a disk-loaded waveguide. In the following description, the term "longitudinal cross section" is used to define the cross section in a plane through the centre axis. The "transverse cross section" is used to define the cross section in a plane orthogonal to the centre axis. A longitudinal centre of an object is halfway down the object's longitudinal axis. For example, the longitudinal centre of a cell is the plane half way along the centre axis of that cell. Each cell has an iris 314 connecting to the preceding cell in the sequence, and an iris 315 connecting to the next cell in the sequence. The irises and cells are centred on the centre axis. In use, the centre axis defines the electron acceleration path, the path along which electrons travel when being accelerated though the acceleration waveguide 300. Generally, cells and irises are axisymmetrical around the centre axis, forming a rounded toroid, i.e. the three-dimensional shape created by sweeping a two-dimensional shape around the axis. Although each of the cells shown in Fig. 3 has the same dimensions, with a fixed radius r and fixed length along the acceleration path 312, some waveguide designs may use coupled cells of varying dimensions, in particular cells of varying length. The length of the cell is chosen to provide a suitable length through which each electron experiences acceleration due to the oscillating RF field. In the waveguide shown in Fig. 3, a "nose cone" 316 is formed on each end of the iris, lengthening the iris along the centre axis to protrude into the cell, for purposes of concentrating the longitudinal electric field component at the beam axis. However, some waveguides do not include a nose cone. The cells are manufactured by welding segments of conductive material together at joining portions. The joining portions of the segments are typically in the longitudinal centre of the cell, marked a. Additionally, the plane at the longitudinal centre of each cell is a plane of symmetry. Many alternative methods may be used to make a cavity, such as brazing and diffusion bonding. One alternative to welding individual cavities together is to instead segment a larger cylindrical piece into multiple cavities. Once a waveguide has been produced, the waveguide may be "tuned" to try to bring each cell to the correct resonance. This can be done by taking a measurement of the electric field created in a waveguide upon the application of radiofrequency energy, introducing a perturbation such as a dent into the waveguide, and then taking another measurement of the electric field to determine whether the resonance is correct. This aligns the optimum path the electrons take and minimises their interaction with the higher order modes. The exemplary acceleration waveguides of Figs. 1 to 3 are conventionally arranged to support a single mode of electron acceleration in order to produce a treatment beam. Disclosed herein is an acceleration waveguide having a geometrical structure arranged to support at least two modes of electron acceleration in order to obtain respective x-rays with respective, different energies. The acceleration waveguide having a geometrical structure arranged to support at least two modes of electron acceleration may, for convenience and brevity of description, be referred to herein as a "multiple mode acceleration waveguide". The multiple mode acceleration waveguide may be part of a linear accelerator for a radiotherapy apparatus, and linear accelerators featuring a multiple mode acceleration waveguide may, for convenience and brevity of description, be referred to herein as "multiple mode" linear accelerators. Such "multiple mode" linear accelerators comprise an acceleration waveguide having a geometrical structure arranged to support at least two modes of electron acceleration, and further comprise a target configured to: generate x-rays for imaging when impacted by electrons accelerated by a first of the at least two modes of electron acceleration, and generate x-rays for radiotherapy when impacted by electrons accelerated by a second of the at least two modes of electron acceleration. The x-rays for radiotherapy may form, or be used to form, a beam of therapeutic radiation, to be used in radiotherapy treatment Fig. 4a shows simulated properties of a treatment mode of an exemplary multiple mode acceleration waveguide 400. Fig. 4b shows simulated properties of an imaging mode of the exemplary multiple mode acceleration waveguide 400 (in this example, the imaging mode is a LOM). In each of Figs. 4a and 4b, a cross-section of an exemplary geometrical structure of the multiple mode acceleration waveguide 400 is shown, a simulated acceleration field 410a, 410b or acceleration gradient of the respective mode of that structure is shown, and a corresponding plot 420a, 420b of the variation of the acceleration gradient with waveguide length Z for that particular mode and structure is shown. The length Z is a distance measured along the acceleration axis of the acceleration waveguide 400. In Fig. 4a, the acceleration waveguide 400 is provided with an RF power input corresponding to a first mode of electron acceleration. The RF power produces or excites, within the geometrical structure of the acceleration waveguide 400, a fundamental mode of electron acceleration having an acceleration field 410a suitable for generating an x-ray treatment beam when accelerated electrons strike an appropriate target. The acceleration field 410a is shown using a gradient fill within cells the acceleration waveguide 400. In Fig. 4b, the acceleration waveguide 400 is provided with an RF input corresponding to the production or excitation of a second, lower order, mode of electron acceleration within the geometrical structure of the acceleration waveguide 400. The resulting acceleration field 410b is shown using a gradient fill within cells of the acceleration waveguide 400. The acceleration field 410b of Fig. 4b is suitable for generating an x-ray imaging beam when accelerated electrons strike an appropriate target. It can be seen that a higher acceleration gradient is provided to electrons for the RF input and corresponding fundamental mode excitation of Fig. 4a than for the RF input and corresponding lower order mode excitation of Fig. 4b. In some examples, the target for producing imaging x-rays and for producing radiotherapy treatment x-rays is the same target. In other examples, a composite target may be used, in which x-rays for imaging may be produced from a particular part of the target that is different to the part of the target from which x-rays for radiotherapy treatment may be produced. The geometrical structure for a multiple mode acceleration waveguide such as the acceleration waveguide 400 of Figs. 4a and 4b may be simulated, designed, and / or evaluated using an approach based on Pareto front optimisation. Such a geometrical structure may be simulated, designed, and / or evaluated by using a Pareto front based approach involving: deciding on the number of treatment modes and the number of imaging modes that would utilise the same target structure, designing a geometrical structure for an acceleration waveguide to utilise a desired frequency based on the chosen RF power source, and including additional Pareto fronts (one for each additional LOM and / or HOM) with additional requirements / restrictions. The additional requirements / restrictions may include: 1) the LOM and / or HOM must be within the operational range / limit of the RF power source, 2) the field distribution of the LOM and / or HOM must maintain the field gradient ratio of the buncher section to be the same as that of the fundamental mode, 3) the resulting field on axis of the LOM and / or HOM will result in the desired deceleration of the beam to achieve the energy spectrum and spot size for either the additional imaging or treatment modality as defined by an idealised Monte Carlo model, 4) a Monte Carlo analysis / simulation based on the beam performance of the LOM and / or HOM and therefore linked to the dose rate from the electromagnetic beam simulation is sufficient to allow the same target to be utilised, and 5) the thermal distribution of the LOM and / or HOM will not result in mechanical destruction of the target. It will be appreciated by those skilled in the art that not all of the steps of the above simulation, design, and / or evaluation approach may be necessary to obtain a particular geometrical structure, and that other design and / or optimisation approaches may be used to identify a suitable geometrical structure, including approaches that do not make use of Pareto fronts. A buncher section of the multiple mode acceleration waveguide may be arranged to support the bunching of electrons for each of at least two modes of electron acceleration. The buncher section typically controls how efficient the linear accelerator is as well as the dose rate generated. Accordingly, in a multiple mode linear accelerator, a buncher should be designed not only for the fundamental mode but also to maintain its characteristics for a HOM and / or LOM to be used for imaging. The multiple mode acceleration waveguide may comprise a buncher section having a geometrical structure simulated, designed, and / or evaluated using the Pareto front based approaches disclosed herein. Such a structure may be particularly advantageous for supporting proper electron bunching behaviour for at least two modes of electron acceleration. In one example, a Pareto front based design / evaluation approach has been used to identify that a particular geometrical structure for an acceleration waveguide fulfils a Pareto design front for a treatment mode, as well as 10 Pareto front requirements for an additional two imaging modes, thereby fulfilling a total of 11 design Pareto fronts. A linear accelerator making use of an acceleration waveguide having the identified geometrical structure allows for a high clinical dose rate with an additional two imaging modes: one a KV mode, and the other an MV mode, both of which can utilise the same target. In the examples of Fig. 4a and Fig. 4b, the acceleration waveguide 400 has a geometrical structure that, when provided with an RF power input at a frequency of about 2.991 GHz, excites a LOM of electron acceleration for producing an imaging beam. The LOM mode has a suitable field in the buncher section to enable electron bunching but, compared to the field of a treatment beam, has a reduced field strength in the relativistic section so that the beam energy is reduced compared to that of a treatment beam. For example, a 45 kV x-ray beam for imaging may be produced from the LOM mode. The 45 kV x-ray beam for imaging may be produced from a 500mA electron injection current. Supplying RF input power at about 2.998 GHz to the acceleration waveguide 400 will excite a fundamental mode and produce a 7 MV x-ray beam for treatment. Accordingly, the geometrical structure of the acceleration waveguide 400 enables alternative imaging and treatment beams to be produced from the same linear accelerator, along the same axis. The geometrical structure of the acceleration waveguide 400 may be evaluated and / or tuned using a Pareto front based simulation approach, such as the Pareto front based approaches disclosed herein. It will be appreciated that, in other examples, alternative geometrical structures may be used in order to support different HOMs and / or LOMs at different frequencies, and that the exact frequencies will depend upon the exact design and / or dimensions of the geometrical structure. Likewise, although in the examples of Figs. 4a and 4b, the x-rays for imaging have an energy of about 45 kV, and the x-rays for radiotherapy have an energy of about 7 MV, in other examples, other energies may be implemented corresponding to the exact design used. Generally, in each example, the x-rays for imaging have an energy in a kV range and the x-rays for radiotherapy have an energy in an MV range. In some examples, the x-rays for imaging have an energy of no less than 35 kV and no more than 150 kV, and the x-rays for radiotherapy have an energy of no less than 1 MV. However, in some examples, the x-rays for imaging from a LOM and / or HOM may have an energy in a MV range, such as no less than 1 MV. Furthermore, although in the examples of Figs. 4a and 4b, the lower order mode is used to generate x-rays for imaging and the fundamental mode is used to generate x-rays for radiotherapy, in other examples, a higher order mode, or combination of lower order mode and higher order mode, may instead be used for imaging. As noted above, a linear accelerator such as that of Fig. 2 comprises an RF power source 214. The RF power source 214 may be a magnetron, a klystron or the like, and is configured to produce and / or amplify RF waves as an output, which are then input into an acceleration waveguide such as the acceleration waveguide 202 of Fig. 2. Advantageously, the modes of the acceleration waveguide 400 of Fig. 4 may each be excited by a suitable magnetron that may be tuned to enable the output of RF power at either about 2.991 GHz or about 2.998 GHz, thereby enabling energy switching in the linear accelerator. In other examples, other sources of RF power may be used. Also disclosed herein is an RF power source for use with the multiple mode linear accelerators of the present disclosure. The RF power source may be configured to provide power of at least two frequencies to the linear accelerator, the at least two frequencies corresponding to respective modes of electron acceleration of the at least two modes of electron acceleration. In particular, the RF power source may be configured to provide power to an acceleration waveguide 400 such as that of Fig. 4. The RF power source may be switchable between the at least two frequencies, such that it alternatively outputs power at one of the at least two frequencies and does not provide power at the least two frequencies simultaneously. The RF power source may be arranged to selectively excite one mode of at least two modes of electron acceleration of the acceleration waveguide. Fig. 5 shows an exemplary configuration of an RF power source 500 for use with a multiple mode acceleration waveguide such as the acceleration waveguide 400 of Fig. 4. In some examples, the RF power source 500 is configurable such that it is capable of alternatively outputting RF waves at two or more frequencies, including at least the frequency which corresponds to the fundamental operating mode of the acceleration waveguide, and at least one other frequency corresponding to a HOM and / or LOM of the acceleration waveguide. The RF power source 500 may be an RF source which can alter one or more of the amplitude, delay and / or pulse width of a pulse in order to alter the RF wave frequency. The RF power source 500 comprises a first RF power generator 524 and a second RF power generator 526, which may each be considered as a RF power 'sub-source'. Each of the first and second RF power generators 524, 526 is configured to produce and / or amplify RF waves, and may each be a magnetron, a klystron or the like. The bandwidth of each of the first and second RF power generators 524, 526 may be such that each first and second RF power generator 524, 526 can produce respective RF waves of a desired frequency. For example, the first RF power generator 524 may produce RF waves suitable for operating a linear accelerator in an imaging mode, and the second RF power generator 526 may produce RF waves suitable for operating a linear accelerator in a radiotherapy treatment mode. In such implementations, solely the first RF power generator 524 may be in operation when the linear accelerator is to be operated in an imaging mode, and solely the second RF power generator 526 may be in operation when the linear accelerator is to be operated in a radiotherapy treatment mode. The first and second RF power generators 524, 526 may be connected and / or coupled using a waveguide junction 528. The waveguide junction 528 is a device used to enable the power in a waveguide to be split, combined or, in some cases, extracted. The waveguide junction 528 can be used to split, combine and / or extract power from the first and second RF power generators 524, 526, such that a number of RF frequency inputs to the multiple mode acceleration waveguide 400 are possible. The waveguide junction 528 may be, for example, a hybrid coupler, and in particular a 'Magic T' (or magic tee) hybrid coupler, which can be used to split or combine the RF inputs from the first and second RF power generators 524, 526. The 'Magic T' comprises four ports, and depending on the coupling formation (i.e. the ports to which each RF power generator is connected), the RF output from the waveguide junction may be the sum or the difference of the individual RF inputs. The waveguide junction 528 is connected to the multiple mode acceleration waveguide 400 via RF transmission apparatus 216. Like in Fig. 2, the RF transmission apparatus 216 is perpendicular to the multiple mode acceleration waveguide 400 central beam axis where it couples the power into the input cell. The RF input connecting pipe or tube is coupled with the acceleration waveguide and joins the acceleration waveguide at a substantially 90° angle. The RF transmission apparatus 216 may include a circulator 218 of any appropriate known type. The RF waves from the RF power source 500 are input into a particular cell of the acceleration waveguide 400 of the multiple mode linear accelerator. The RF transmission apparatus 216 that connects the waveguide junction 528 to the input cell of the multiple mode acceleration waveguide 400 may comprise a waveguide network and may contain an RF window which may separate a vacuum envelope from an SF6 envelope. The multiple mode linear accelerators disclosed herein may further comprise a controller 570. The controller 570 may form part of the control system of the radiotherapy apparatus (such as in the radiotherapy device 100 of Fig. 1 or the radiotherapy system 700 of Fig. 7) or may be entirely separate from the control system. The controller 570 may comprise a computer-based system which may be used for controlling or operating various parts of the systems, devices, methods and apparatuses disclosed herein. The computer-based system can be implemented in software, firmware and / or hardware and may comprise a computer-readable medium containing instructions that, when executed by a processor, cause the system to perform any of the methods described herein. The controller 570 is configured to control features of the multiple mode linear accelerator. In particular, the controller 570 is configured, for each of at least two modes of electron acceleration, to maintain an electromagnetic field through the buncher section of the multiple mode acceleration waveguide of the linear accelerator. The electromagnetic field may be controlled and / or maintained by controlling the RF being applied to the acceleration waveguide. The controller is thus arranged to maintain and / or selectively maintain an electromagnetic field for bunching electrons for acceleration by one of at least two modes of electron acceleration. The controller 570 thus enables appropriate electron bunching to be achieved in the buncher section regardless of the mode of operation of the acceleration waveguide, and acts to prevent electron bunching from breaking down for a particular mode of the multiple modes. The buncher can therefore function at both the fundamental operating mode and the at least one HOM and / or LOM. Thus, the controller 570 is configured to operate the buncher in a way that allows the selection of a particular mode of acceleration and thus a corresponding imaging or treatment energy to be produced by the linear accelerator. In some examples, the controller 570 may control the RF power source 500 in order to control switching of the frequency of the RF power source 500 and resultant switching of the energy of the beam of the linear accelerator. A distinct advantage of this type of energy switching is the potential speed at which the treatment mode can be switched between a treatment mode and an imaging mode. Mechanical methods of switching are limited by motor speeds to switching times in the order of milliseconds or seconds. The present disclosure enables switching times in the order of microseconds to milliseconds if using, for example, a fast tuned magnetron as an RF power source. Hardware components of controller 570 may include one or more computers (e.g., general purpose computers, workstations, servers, terminals, portable / mobile devices, etc.); processors (e.g., central processing units (CPUs), graphics processing units (GPUs), microprocessors, digital signal processors (DSPs), field programmable gate arrays (FPGAs), special-purpose or specially-designed processors, etc.); memory / storage devices such as a memory (e.g., read-only memories (ROMs), random access memories (RAMs), flash memories, hard drives, optical disks, solid-state drives (SSDs), etc.); input devices (e.g., keyboards, mice, touch screens, mics, buttons, knobs, trackballs, levers, handles, joysticks, etc.); output devices (e.g., displays, printers, speakers, vibration devices, etc.); circuitries; printed circuit boards (PCBs); or other suitable hardware. Software components of controller 570 may include operation device software, application software, etc. Fig. 6 is a flowchart diagram showing a method 600 of operation of a multiple mode linear accelerator such as those discussed in relation to Figs. 4a, 4b, and 5. The method 600 comprises, at block 640, providing RF power to a multiple mode linear accelerator to accelerate electrons by one of at least two modes of electron acceleration of the multiple mode acceleration waveguide. The RF power may be provided by any of the RF power sources disclosed herein, such as the RF power source 500 of Fig. 5 and / or the tunable magnetron example of Figs. 4a and 4b. Optionally, the method 600 may also comprise, at block 610, selecting an RF power to output from the RF power source to the acceleration waveguide and / or the linear accelerator. At block 610, an RF power is chosen based on the desired beam energy. This may be the beam energy needed to operate the linear accelerator in a radiotherapy treatment mode, or may be the beam energy needed to operate the linear accelerator in an imaging mode. The RF power source may comprise a single RF power source, such as a magnetron or a Klystron, in which case the RF power source is then tuned to the required frequency. Alternatively, the RF power input may be a configurable RF power source and one or more of the amplitude, delay and / or pulse width of a pulse may be altered in order to achieve the desired RF frequency output. In examples where the RF power source comprises more than one RF power sub-source, the coupling formation of the ports of the waveguide junction and the individual RF power sub-sources is arranged such that the desired RF frequency is output from the waveguide junction. Different coupling formations are required in order to split, combine or extract the RF frequency, and as such the frequency output will depend on the coupling formation for a given example. The method 600 further comprises, at block 620, providing electrons to the acceleration waveguide. A beam of electrons is generated by the source of electrons, which may be, for example, an electron gun. The electrons are injected into the acceleration waveguide and accelerated along the central axis of the waveguide. The injection of electrons into the acceleration waveguide may be synchronised with the pulsing of the radiofrequency waves into the acceleration waveguide. The beam of electrons enter the buncher section of the waveguide where the phase of the RF wave generated at block 610 decelerates some electrons to allow slower electrons to catch up, concentrating the electrons in bunches. The electrons are then free to move together in so called "packets" or "bunches", and the bunches quickly accelerate to relativistic speeds through the subsequent cells of the acceleration waveguide. The method 600 optionally further comprises, at block 630, using a controller to control and / or maintain an electromagnetic field in the buncher section of the waveguide such that bunching of electrons occurs. The bunched electrons are then accelerated by the respective mode of the multiple modes of the acceleration waveguide. As described herein, the controller may be selectively switchable to control the RF to maintain the electromagnetic field in a manner that corresponds to the desired mode of electron acceleration, such that the buncher section may support bunching of electrons for any of the modes of electron acceleration supported by the geometrical structure of the acceleration waveguide. The approaches disclosed herein may be applied to a traveling wave linear accelerator or to a standing wave linear accelerator. Referring to the apparatuses disclosed herein, variations in design and components are possible or desirable depending upon the application requirements. For example, requirements may vary depending upon the desired type and energy of the treatment beam or depending upon the mechanical or structural design of the overall device in which the apparatus is to be used, such as the device 100 of Fig. 1. In some implementations, the RF power source 214 may be a klystron, rather than a magnetron. Similarly, in some implementations, the RF power source 214 may be operated continuously rather than in a pulsed manner. The beam generation system 200 also comprises several other components and systems as will be understood by the skilled person. For example, in order to ensure the linac or beam generation system does not leak radiation, appropriate shielding is also provided. The whole system is cooled by a water cooling system (not shown in the figures). The water cooling system may be used, in particular, to cool the acceleration waveguide 202, the target 208, and the RF power source 214. A computer-based system may be used for controlling or operating various parts of the systems, devices, methods and apparatuses disclosed herein. The computer-based system can be implemented in software, firmware and / or hardware and may comprise a computer-readable medium containing instructions that, when executed by a processor, cause the system to perform any of the methods described herein. Fig. 7 illustrates a block diagram of one implementation of a radiotherapy system 700. The radiotherapy system 700 comprises a computing system 710 within which a set of instructions, for causing the computing system 710 to perform any one or more of the methods discussed herein, may be executed. The computing system 710 shall be taken to include any number or collection of machines, e.g. computing device(s), that individually or jointly execute a set (or multiple sets) of instructions to perform any one or more of the methods discussed herein. That is, hardware and / or software may be provided in a single computing device, or distributed across a plurality of computing devices in the computing system. In some implementations, one or more elements of the computing system may be connected (e.g., networked) to other machines, for example in a Local Area Network (LAN), an intranet, an extranet, or the Internet. One or more elements of the computing system may operate in the capacity of a server or a client machine in a client-server network environment, or as a peer machine in a peer-to-peer (or distributed) network environment. One or more elements of the computing system may be a personal computer (PC), a tablet computer, a set-top box (STB), a Personal Digital Assistant (PDA), a cellular telephone, a web appliance, a server, a network router, switch or bridge, or any machine capable of executing a set of instructions (sequential or otherwise) that specify actions to be taken by that machine. The computing system 710 includes controller circuitry 711 and a memory 713 (e.g., read-only memory (ROM), flash memory, dynamic random access memory (DRAM) such as synchronous DRAM (SDRAM) or Rambus DRAM (RDRAM), etc.). The memory 713 may comprise a static memory (e.g., flash memory, static random access memory (SRAM), etc.), and / or a secondary memory (e.g., a data storage device), which communicate with each other via a bus (not shown). Controller circuitry 711 represents one or more general-purpose processors such as a microprocessor, central processing unit, accelerated processing units, or the like. More particularly, the controller circuitry 711 may comprise a complex instruction set computing (CISC) microprocessor, reduced instruction set computing (RISC) microprocessor, very long instruction word (VLIW) microprocessor, processor implementing other instruction sets, or processors implementing a combination of instruction sets. Controller circuitry 711 may also include one or more special-purpose processing devices such as an application specific integrated circuit (ASIC), a field programmable gate array (FPGA), a digital signal processor (DSP), network processor, or the like. One or more processors of the controller circuitry may have a multicore design. Controller circuitry 711 is configured to execute the processing logic for performing the operations and steps discussed herein. The computing system 710 may further include a network interface circuitry 718. The computing system 710 may be communicatively coupled to an input device 720 and / or an output device 730, via input / output circuitry 717. In some implementations, the input device 720 and / or the output device 730 may be elements of the computing system 710. The input device 720 may include an alphanumeric input device (e.g., a keyboard or touchscreen), a cursor control device (e.g., a mouse or touchscreen), an audio device such as a microphone, and / or a haptic input device. The output device 730 may include an audio device such as a speaker, a video display unit (e.g., a liquid crystal display (LCD) or a cathode ray tube (CRT)), and / or a haptic output device. In some implementations, the input device 720 and the output device 730 may be provided as a single device, or as separate devices. In some implementations, the computing system 710 may comprise image processing circuitry 719. Image processing circuitry 719 may be configured to process image data 780 (e.g. images, or imaging data), such as medical images obtained from one or more imaging data sources, a treatment device 750 and / or an image acquisition device 740. Image processing circuitry 719 may be configured to process, or pre-process, image data. For example, image processing circuitry 719 may convert received image data into a particular format, size, resolution or the like. In some implementations, image processing circuitry 719 may be combined with controller circuitry 711. In some implementations, the radiotherapy system 700 may further comprise an image acquisition device 740 and / or a treatment device 750, such as those disclosed herein in the example of Fig. 1. The image acquisition device 740 and the treatment device 750 may be provided as a single device. In some implementations, treatment device 750 is configured to perform imaging, for example in addition to providing treatment and / or during treatment. The treatment device 750 comprises the main radiation delivery components of the radiotherapy system, such as the beam generation systems and linear accelerator components disclosed herein. Image acquisition device 740 may be configured to perform positron emission tomography (PET), computed tomography (CT), and magnetic resonance imaging (MRI). Image acquisition device 740 may be configured to output image data 780, which may be accessed by computing system 710. Treatment device 750 may be configured to output treatment data 760, which may be accessed by computing system 710. Computing system 710 may be configured to access or obtain treatment data 760, planning data 770 and / or image data 780. Treatment data 760 may be obtained from an internal data source (e.g. from memory 713) or from an external data source, such as treatment device 750 or an external database. Planning data 770 may be obtained from memory 713 and / or from an external source, such as a planning database. Planning data 770 may comprise information obtained from one or more of the image acquisition device 740 and the treatment device 750. The various methods described above may be implemented by a computer program. The computer program may include computer code (e.g. instructions) 810 arranged to instruct a computer to perform the functions of one or more of the various methods described above. The steps of the methods described above may be performed in any suitable order. For example, step 620 of method 600 may be performed before, after, simultaneously or substantially simultaneously with step 610. The computer program and / or the code 810 for performing such methods may be provided to an apparatus, such as a computer, on one or more computer readable media or, more generally, a computer program product 800)), depicted in Fig. 8. The computer readable media may be transitory or non-transitory. The one or more computer readable media 800 could be, for example, an electronic, magnetic, optical, electromagnetic, infrared, or semiconductor system, or a propagation medium for data transmission, for example for downloading the code over the Internet. Alternatively, the one or more computer readable media could take the form of one or more physical computer readable media such as semiconductor or solid state memory, magnetic tape, a removable computer diskette, a random access memory (RAM), a read-only memory (ROM), a rigid magnetic disc, and an optical disk, such as a CD-ROM, CD-R / W or DVD. The instructions 810 may also reside, completely or at least partially, within the memory 713 and / or within the controller circuitry 711 during execution thereof by the computing system 710, the memory 713 and the controller circuitry 711 also constituting computer-readable storage media. In an implementation, the modules, components and other features described herein can be implemented as discrete components or integrated in the functionality of hardware components such as ASICS, FPGAs, DSPs or similar devices. A "hardware component" is a tangible (e.g., non-transitory) physical component (e.g., a set of one or more processors) capable of performing certain operations and may be configured or arranged in a certain physical manner. A hardware component may include dedicated circuitry or logic that is permanently configured to perform certain operations. A hardware component may comprise a 5 special-purpose processor, such as an FPGA or an ASIC. A hardware component may also include programmable logic or circuitry that is temporarily configured by software to perform certain operations. In addition, the modules and components can be implemented as firmware or functional circuitry 10 within hardware devices. Further, the modules and components can be implemented in any combination of hardware devices and software components, or only in software (e.g., code stored or otherwise embodied in a machine-readable medium or in a transmission medium). Those skilled in the art will recognise that a wide variety of modifications, alterations, and 15 combinations can be made with respect to the above described examples without departing from the scope of the disclosed concepts, and that such modifications, alterations, and combinations are to be viewed as being within the ambit of the disclosed concepts. Those skilled in the art will also recognise that the scope of the invention is not limited by the examples described herein but is instead defined by the appended claims.

Claims

1. A linear accelerator for a radiotherapy apparatus, the linear accelerator comprising:an acceleration waveguide having a geometrical structure arranged to support at least two modes of electron acceleration; anda target configured to:generate x-rays for radiotherapy when impacted by electrons accelerated by a first of the at least two modes of electron acceleration, andgenerate x-rays for imaging when impacted by electrons accelerated by a second of the at least two modes of electron acceleration.

2. The linear accelerator of claim 1, wherein the x-rays for imaging have an energy of no less than 35 kV and no more than 150 kV, and the x-rays for radiotherapy have an energy of no less than 1 MV.

3. The linear accelerator of any preceding claim, wherein the x-rays for imaging have an energy of about 45 kV, and the x-rays for radiotherapy have an energy of about 7 MV.

4. The linear accelerator of any preceding claim, wherein the at least two modes of electron acceleration comprise a lower order mode for generating x-rays for imaging and a fundamental mode for generating x-rays for radiotherapy.

5. The linear accelerator of any preceding claim, wherein the second of the at least two modes of electron acceleration is a lower order mode.

6. The linear accelerator of any preceding claim, wherein the second of the at least two modes of electron acceleration corresponds to an input power frequency of about 2.991 GHz.

7. The linear accelerator of any preceding claim, wherein the first of the at least two modes of electron acceleration corresponds to an input power frequency of about 2.998 GHz.

8. The linear accelerator of any preceding claim, wherein the acceleration waveguide comprises a buncher section arranged to support the bunching of electrons for each of the at least two modes of electron acceleration.

9. The linear accelerator of any preceding claim, further comprising a controller arranged to maintain an electromagnetic field for bunching electrons for acceleration by one of the at least two modes of electron acceleration.

10. The linear accelerator of any preceding claim, further comprising a radio frequency power source configured to provide power of at least two frequencies to the linear accelerator, the at least two frequencies corresponding to respective modes of electron acceleration of the at least two modes of electron acceleration.

11. The linear accelerator of claim 10, wherein the radio frequency power source is arranged to be switchable between the at least two frequencies.

12. The linear accelerator of claim 10 or claim 11, wherein the radio frequency power source comprises a first RF power generator and a second RF power generator.

13. The linear accelerator of claim 12, wherein the first RF power generator and the second RF power generator are coupled using a hybrid coupler.

14. A radio frequency power source for use with the linear accelerator of any preceding claim, wherein the radio frequency power source is configured to provide power of at least two frequencies to the linear accelerator, the at least two frequencies corresponding to respective modes of electron acceleration of the at least two modes of electron acceleration.

15. The radio frequency power source of claim 14, wherein the radio frequency power source is arranged to be switchable between the at least two frequencies.

16. The radio frequency power source of any of claims 14 to 15, wherein the radio frequency power source comprises a first RF power generator and a second RF power generator.

17. The radio frequency power source of claim 16, wherein the first RF power generator and the second RF power generator are coupled using a hybrid coupler.

18. A method of operating the linear accelerator of any of claims 1 to 13, the method comprising:providing electrons to the acceleration waveguide of the linear accelerator; and providing radio frequency power to the linear accelerator to accelerate electrons by one of the at least two modes of electron acceleration.

19. The method of claim 18, the method further comprising selecting the radio frequency power using a radio frequency power source configured to provide power of at least two frequencies to the linear accelerator, the at least two frequencies corresponding to respective modes of electron acceleration of the at least two modes of electron acceleration.

20. The method of claim 19, wherein the radio frequency power source is arranged to be switchable between the at least two frequencies.

21. The method of claim 19 or claim 20, wherein the radio frequency power source comprises a first RF power generator and a second RF power generator.

22. The method of claim 21, wherein the first RF power generator and the second RF power generator are coupled using a hybrid coupler.

23. The method of any of claims 18 to 22, the method further comprising controlling an electromagnetic field in a buncher section of the acceleration waveguide to maintain the electromagnetic field for bunching electrons for acceleration by one of the at least two modes of electron acceleration.

24. One or more computer readable media comprising instructions, that, when executed by one or more processors, cause the one or more processors to perform the method of any of claims 18 to 23.

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