WO2009050626A1

WO2009050626A1 - Imaging system with distributed sources and detectors

Info

Publication number: WO2009050626A1
Application number: PCT/IB2008/054187
Authority: WO
Inventors: Hermann Schomberg; Randall Peter Luhta; Rainer Pietig
Original assignee: Koninklijke Philips Electronics N.V.; Philips Intellectual Property & Standards Gmbh
Priority date: 2007-10-19
Filing date: 2008-10-13
Publication date: 2009-04-23

Abstract

The invention relates to a cone beam computed tomography imaging system comprising a plurality of distributed radiation sources (20) arranged to emit radiation beams from a plurality of emission locations and a plurality of distributed radiation detector modules (21). The pluralities of sources and detector modules are arranged along source trajectories and detector trajectories, respectively, such that a straight line that starts at a point on the source trajectory and passes through the isocenter of the system intersects the detector trajectory. A radiation beam is not significantly blocked by a portion of a radiation detector module between the emission location and the volume of interest. The system is not hampered from self-blocking.

Description

IMAGING SYSTEM WITH DISTRIBUTED SOURCES AND DETECTORS

FIELD OF THE INVENTION

The invention relates to an x-ray computed tomography (CT) system. Moreover, the invention relates to a method of operating a CT system and to a computer program product.

BACKGROUND OF THE INVENTION

Imaging systems, and in particular x-ray imaging systems, are utilized for various applications in both medical and non-medical fields. Medical x-ray imaging systems such as computed tomography (CT) systems are capable of producing exact cross- sectional or volumetric image data that express certain physical properties related to the human or animal body. Reconstruction of three-dimensional (3D) images representing a volume of interest has been applied in the medical field for some time.

Cardiac volume imaging is an important application of CT. Cardiac volume imaging is typically done with helical CT and retrospective gating. Even though this approach works reasonably well, it is hampered by limited temporal resolution and high radiation dose.

In helical CT, the patient table is translated during the scan while the gantry of the CT scanner rotates. Once multi-row detectors become available that are, for instance, 15 cm "high", the need for translating the patient table disappears. But then the rotation speed of the gantry becomes a limiting factor. The time window for collecting the data for a 3D snapshot image of the beating heart during diastole is only approximately 100 ms long, which means that the gantry has to complete one rotation in about 180 ms. Such a gantry would be very costly.

The ability of collecting enough data for an image within the required short time is offered by Electron Beam CT (EBCT). An EBCT scanner has a stationary detector and a stationary, huge, scanning electron beam x-ray tube. However, the scanning electron beam x-ray tube remains a drawback of EBCT; it is bulky and costly, and its x-ray power is limited. In US 2007/0009088 a system is disclosed which seeks to reduce challenges associated with movement of a source and/or a detector. In the disclosure, an x-ray imaging system is provided which includes a distributed x-ray source configured to emit x-rays from a plurality of emission points and a detector. Embodiments of the geometric arrangements of the emission points and the detector are provided. However, with the disclosed geometric arrangements, when they are used without mechanical movement, it is not possible to collect a set of cone beam projections from which one can reconstruct a high-quality CT image.

The inventors of the present invention have appreciated that an improved imaging system is of benefit, and have in consequence devised the present invention.

SUMMARY OF THE INVENTION

The invention preferably seeks to mitigate, alleviate or eliminate one or more of the above mentioned disadvantages singly or in any combination. In particular, it may be seen as an object of the present invention to provide an imaging system that solves the above mentioned problems, or other problems, of the prior art related to imaging systems.

This object and several other objects are obtained in a first aspect of the invention by providing a computed tomography imaging system comprising

- a plurality of distributed radiation sources arranged to emit radiation beams from a plurality of emission locations, the plurality of emission locations being arranged substantially along a source trajectory, the radiation penetrating a volume of interest around an isocenter;

- a plurality of distributed radiation detector modules, the plurality of radiation detector modules being arranged substantially along a detector trajectory; wherein the source and detector trajectories are such that a straight line that starts at a point on the source trajectory and passes through the isocenter intersects the detector trajectory; and wherein the radiation beam that is emitted from an emission location on the source trajectory is not significantly blocked by a portion of a radiation detector module between the emission location and the volume of interest. The isocenter is a distinguished spatial point conceptually attached to the system.

The invention has the insight that in order to provide a system that is not hampered from self- blocking, a special arrangement of both the distributed radiation sources and of the distributed detector modules is needed. A self-blocking system suffers from a blocking of the emitted radiation by intervening portions of the detector. It is an advantage of the present invention that an imaging system may be provided which is not self-blocking, where the plurality of distributed radiation sources may be configured to remain stationary with respect to the volume of interest, and where the plurality of distributed radiation detector modules may be configured to remain stationary with respect to the volume of interest. A CT scanner without moving parts may thereby be provided.

It is another advantage of an imaging system according to the first aspect of the present invention that the source trajectory is complete with respect to a sizeable volume of interest around the isocenter. In the context of the present disclosure, a source trajectory is said to be complete with respect to a volume V, if every plane that intersects V also intersects the source trajectory. This condition is also known as Tuy's completeness condition. A planar source trajectory cannot be complete with respect to a true volume. If a source trajectory is complete with respect to V, then an accurate, three-dimensional image of the content of V may be reconstructed from the cone beam projections of V taken along the source trajectory, provided these cone beam projections are not truncated.

In an exemplary embodiment, the plurality of distributed radiation sources ranges from 500 to 1000 radiation sources. It may be advantageous to use a number of sources in this range in order to achieve a sufficient resolution of the reconstructed image.

In an exemplary embodiment, the radiation sources are carbon-nanotube-based x-ray sources. It is an advantage to use such a source type, since such sources may be fabricated suitably small so that a large number of sources may be placed along the source trajectory. The total detector area resembles a non-planar, specially curved strip, or sheet. In an exemplary embodiment, the detector is built from identical detector modules, the odd and even numbered detector modules may be arranged alternately with a first and second distance from the isocenter of the volume of interest. A closer packing of the detector modules is thereby achieved.

In an exemplary embodiment, the plurality of distributed radiation sources and the plurality of distributed detector modules are individually addressable. Advantageous scan protocols which avoid overheating of the sources are thereby rendered possible.

In a second aspect of the invention a method of operating a cone beam CT imaging system is provided.

The method of operation may operate a imaging system in accordance with the first aspect of the invention, where the at least first radiation source and the at least first radiation detector module are operated such that radiation emitted from an at least first radiation source passes through the volume of interest and is detected by an at least first radiation detector module.

In exemplary embodiments, groups of radiation detector modules are associated with the radiation sources so that for a first radiation source in the plurality of the radiation sources, and for at least a second radiation source in the plurality of radiation sources, the associated first and the at least second group of radiation detectors do not overlap. This is advantageous, since radiation source activation sequences and detector readout sequences may be provided, which provide fast scanning without suffering form overheating of the source.

In an exemplary embodiment, the sequence in which the radiation sources are turned on, is correlated to an input signal. A scan protocol which allows for long scan sequences of dynamic organs, such as the heart, may thereby be provided.

In a third aspect of the invention, a computer program product having a set of instructions, when in use on a computer, to cause the computer to perform the method of the second aspect. Such computer program product may be implemented in accordance with a system of the first aspect of the invention.

In general, the various aspects of the invention may be combined and coupled in any way possible within the scope of the invention. These and other aspects, features and/or advantages of the invention will be apparent from and elucidated with reference to the embodiments described hereinafter.

BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the invention will be described, by way of example only, with reference to the drawings, in which

FIG 1 shows a cross- sectional view of a schematic exemplary embodiment of a cone beam computed tomography apparatus;

FIG. 2 shows a simplified perspective view of a schematic exemplary embodiment of the radiation source arrangement and detector arrangement;

FIG. 3 illustrates graphs of the functions a(λ), β(λ);

FIG. 4 illustrates an exemplary embodiment of a source module;

FIG. 5 illustrates exemplary embodiments of detector modules and the arrangement of a source;

FIG. 6 schematically illustrates a front view of the source/detector arrangement of an example scanner; and

FIG. 7 illustrates a flow diagram of operating an imaging system in accordance with embodiments of the present invention. DESCRIPTION OF EMBODIMENTS

Embodiments of the present invention are disclosed in connection with an x-ray cone beam computed tomography (CT) scanner or imaging system. Specific numerical values are provided below in connection with example embodiments, these numerical values are provided in connection with an example scanner. It is however clear to the skilled person that numerical values provided in connection with the example scanner, are only provided as example values, and do not limit the scope of the invention.

FIG 1 shows a cross- sectional view of a schematic exemplary embodiment of a cone beam computed tomography apparatus. The apparatus comprises a radiation source 1 capable of emitting x-ray radiation in the form a cone beam 2. The apparatus is associated with a center point called the isocenter 3. Moreover, the apparatus comprises a radiation detector arrangement 4 which is opposite to the radiation source. The apparatus is intended to acquire transversely non-truncated cone beam projections of a certain volume of interest 5 around the isocenter 3. The volume of interest is large enough to contain a human heart, for example. The sensitive area 6 of the detector faces the radiation source. A straight line 8 starting at the center point of the sensitive area and passing through the isocenter meets the center of the emission location 7 of the radiation source.

FIG. 2 shows a simplified perspective view of a schematic exemplary embodiment of the radiation source arrangement and detector arrangement, as well as the arrangement of a collimator slit. For clarity reasons, the detector modules are indicated by brick-shaped cuboids arranged along a detector trajectory 21, whereas locations of the radiation sources are arranged along a source trajectory 20, depicted by a dotted line, and the center line of the collimator slit 22 is depicted by straight lines. The trajectories defined by the source arrangement, the detector arrangement and the collimator arrangement each extend on the surfaces of isocentric spheres with different radii and each resembles a portion of the boundary curve of a saddle. The diameter of the sphere on which the source trajectory is located is slightly larger than the diameter of the sphere on which the detector trajectory is located. These spheres are also referred to as source spheres and detector spheres respectively. In an apparatus in accordance with exemplary embodiments of the present invention, a plurality of distributed radiation sources are arranged to emit radiation beams from a plurality of emission locations being arranged substantially along the source trajectory 20, and a plurality of distributed radiation detector modules are arranged substantially along a detector trajectory 21. The plurality of distributed radiation sources and the plurality of distributed radiation detector modules are configured to remain stationary with respect to the volume of interest. The source and detector trajectories are such that a straight line that starts at a point on the source trajectory and passes through the isocenter intersects the detector trajectory. An important aspect of this geometry is that the radiation beam that is emitted from an emission location on the source trajectory is not significantly blocked by a portion of a radiation detector between the emission location and the volume of interest.

The collimator 22 represents a slit whose center line extends essentially parallel to the source trajectory. Thus, the center line of the collimator slit is essentially a scaled image of the source trajectory.

The plurality of radiation sources may be in the form of small x-ray sources that are aligned along the source trajectory. In exemplary embodiments, the number of sources will lie in the range of 500 and 1000, such as in the range of 600 to 800. The sources may be grouped into identically built modules, e.g. with 9 orlO adjacent sources per module, an example embodiment of source modules is disclosed below in connection with FIG. 4. The plurality of detector modules may be arranged in a detector strip, which is assembled from identically built detector modules, where the sensitive area of each module is a rectangular, possibly tiled, array of small detector elements. Embodiments of the detector modules are disclosed below in connection with FIG. 5. The collimator slit confines the raw cone beams of x-rays that emanate from the sources such that the surviving x-rays hit the desired opposite portion of the detector strip.

In the following, the source trajectory, the detector trajectory and the center line of the collimator slit are specified formally in terms of parametric representations. To this end, a

Cartesian x-y-z coordinate system 23 is attached to the isocenter of the scanner such that the y- axis points horizontally towards the foot end of the patient table, the z-axis points vertically upwards, and the x-axis points horizontally to the right. In FIG. 2, the origin of the coordinate system 23 has been displaced from the isocenter for clarity reasons. The coordinate system 23 thus merely indicates the x-y-z directions. In this coordinate system, the source trajectory is given by

B(λ) = -r_s (cosa(λ)cos β(λ),ύn β(λ) -ύn a(λ)cos β(λ)) (1)

where r_s is the distance from the isocenter to the source trajectory and where a(λ), β(λ) are functions of an abstract parameter λ, with λ = 0 corresponding to the beginning of the source trajectory and λ = 1 corresponding to the end of the source trajectory. In an exemplary embodiment, r_s = 900 mm.

In an exemplary embodiment, the sensitive area of the odd and even numbered detector modules are arranged alternately with a first and second distance from the isocenter, where the center points of the sensitive areas of the odd-numbered modules lie on a first detector trajectory, and the center points of the sensitive areas of the even-numbered modules lie on a second detector trajectory that is similar to the first curve, but has a slightly larger radius. The two detector curves lie on isocentric spheres, too. The long axis of each detector module is perpendicular to the module's detector trajectory and tangential to its detector sphere. Adjacent detector modules are tightly packed side by side. As a result of this arrangement, an observer looking from anywhere on the source curve towards the opposite portion of the detector strip would see virtually no gap between the sensitive areas of adjacent detector modules.

Similar to the source trajectory, the detector trajectory for the odd and even numbered detector modules are given by

E₁ (λ) = r_dl (cos a(λ) cos β(λ), sin β(λ) - sin a(λ) cos β(λ)) (2) b₂ (λ) = r_d2 (cos a(λ) cos β(λ), sin β(λ) - sin a(λ) cos β(λ)) (3) In an exemplary embodiment, λ is taken in the range: -0.4 < λ < 1.04, and r_d\ = 675 mm and r_d2 = 700 mm.

The collimator trajectory is given by

c (λ) = -r_c (cos a(λ) cos β(λ), sin β(λ\- sin a(λ) cos β(λ)) (4)

In an exemplary embodiment, λ is taken in the range: -0.4 < λ < 1.04, and with r_c = 820 mm.

In the source/detector arrangement shown in FIG.2, the radius r_s of the source sphere is greater than the radii ra_\ and rai of the detector spheres. It is also possible to make the radius r_s smaller than the radii ra\ and rai.

FIG. 3 illustrates a(λ), β(λ), in the form of (180/π)α(/l) and (180/π)y?(/t) in order to convert radians to degrees. FIG. 3A shows a(λ) (denoted with reference numeral 30), whereas FIG.

3B shows β(λ) (denoted with reference numeral 31). The function a(λ) is in the form of a straight line, whereas the function β(λ) can not be expressed by a simple closed formula. The geometry of the source trajectory, the detector trajectory, and the collimator trajectory, and more specifically, the functions a(λ) and β(λ) are further disclosed in the published international patent application WO 2005/104952, especially in connection with FIGS. 3-6 and 8-13. This disclosure is hereby incorporated by reference.

FIG. 4 illustrates an exemplary embodiment of a source module, wherein the radiation sources are carbon-nanotube-based x-ray sources, e.g. by using carbon nanotubes as cold electron emitters

In an exemplary embodiment, the long and curvilinear arrangement of sources may be assembled from short, identically built modules 49 with a rectilinear arrangement of sources. Each module may contain a small number of sources, for example 9 or 10. Suitable modules can be built, for example, by modifying the source modules as described in by J. Zhang et al., in the article: "Stationary scanning x-ray source based on carbon nanotube field emitters," Applied Physics Letters 86, 184104, 2005. This article is hereby incorporated by reference. FIG. 4A illustrates a schematically view of a source module in a cross- sectional front-view. FIG. 4B illustrates the source module in cross- sectional side-view. A number (here 9) of cold electron emitters 40 based on carbon nanotubes are placed on the cathode 41. The electron beams 48 are emitted towards the anode 42. The beam is shaped and focused by electrode gates 43. The electron beam is focused on the tilted anode surface 44, resulting in the emission of x-rays 45. The x-rays leave the source module through an x-ray window 46 which limits the transversal cone angle of the raw cone beam. The axial shaping of the cone beams is done by the collimator slit 22 shown in FIG. 2. The anodes of the x-ray sources may actively be cooled by using water or alternative cooling means 47. The cooling agent may, e.g. pass through a hollow anode 400.

The source modules may be mounted onto a properly curved rail, such that their focal spots are aligned essentially along the desired source trajectory. The precise locations of the focal spots can be measured and is favorably used for the reconstruction. The modules may be tightly packed side by side. Since the source curve is bent, there will be small wedge-shaped gaps between adjacent source modules. These gaps cause no harm.

The source trajectory of the example scanner is about L_s = 3600 mm long. In this case, the distance between equidistantly distributed source points is d_s = 4.5 mm for 800 sources and d_s

= 6 mm for 600 sources.

FIG. 5 illustrates exemplary embodiments of detector modules and the arrangement of the source 50, the isocenter 51 and the detector 54, 55. In the exemplary embodiments, the detector strip is assembled from smaller, flat, identically built detector modules 56, 57. Each detector module may in an exemplary embodiment in itself be a rectangular, planar detector that is organized as a 2D array of small detector elements or tiles. The example modules 54, 56 as shown FIG. 5A are brick-shaped and have a sensitive area of 40 mm by 280 mm. The detector modules may be mounted onto a properly curved rail. The precise locations of the active detector areas can be measured and is favorably used during the reconstruction. Since the desired detector strip is not cylindrical, there would remain small wedge-shaped gaps between adjacent detector modules, when the modules would be tightly packed side by side along one of the detector curves. However, when tight packing is combined with alternating radial distances, as indicated by the radii 52 and 53, as shown in FIG. 5A it can be arranged that an observer positioned on a source location and looking towards the opposite portion of the detector strip sees virtually no gaps between the sensitive areas of adjacent modules. The modules can be packed even tighter when they are given a T-shaped cross-section 55, 57, as illustrated in FIG. 5B. A T-shaped cross-section may therefore in embodiments be preferred over a rectangular cross-section.

As a general remark, a source-focused anti-scatter grid cannot be used, as with all CT scanners that vary the source position relative to the detector during a scan.

FIG. 6 schematically illustrates a front view of the source/detector arrangement of the example scanner. Again for clarity reasons, only the detector modules 60 are shown as cuboids, whereas the sources modules 61 are shown as dots. The collimator slit is shown as a solid line. Scan protocols are disclosed in connection with FIG. 6.

In an exemplary embodiment, the sources 61 are numbered consecutively from 1 (marked by reference numeral 62) through N_s (marked by reference numeral 63). Similarly, the detector modules 60 are numbered consecutively from 1 (marked by reference numeral 64) through N_d (marked by reference numeral 65). In the exemplary embodiment, the plurality of distributed radiation sources and the plurality of distributed detector modules are individually addressable and may be individually operated in the sense that they may be turned on, turned off, read out, have specific operation parameters set, etc. In exemplary embodiments, the detector modules are build from detector elements. In such embodiments, the elements may be individually addressable, however at least for some applications, all elements of a module may be individually addressable as a group.

For each radiation source of the plurality of radiation sources, a group of radiation detectors may be associated to the radiation source so as to detect at least a part of the emitted radiation beam. When the zth source fires, i.e. is turned on (as an example, the zth source may be taken as the source marked by reference numeral 66), it emits a "raw" cone beam 66' of x-rays towards the opposite portion of the detector strip. To measure a cone beam projection with the cone beam starting at the zth source, a group of detector modules 67 opposite to the source is activated and read out. For brevity, this group of detector modules will be called the zth detector group 67. The x-rays that start at the zth source and hit the zth detector group define the zth cone beam proper (as opposed to the raw cone beam that leaves the zth source), or simply the zth cone beam. For each i, there are two numbers, Ji₀(Z) and Jh₁(Z), such that the detector modules of the zth detector group are numbered, Ji₀(O, JIo(O⁺I, •••, Jn(O- As i increases, the numbers J₀(O and Jh₁(O remain constant for a small range of z's, and then increase by one, etc. The number of detector modules of each detector group is chosen sufficiently large so that the transversal cone angle of the associated cone beam is about 50°; as a result, the cone beam projections of a non-obese patient will normally not be transversely truncated. By an appropriate design of the x-ray window of each source one can tailor the transversal cone angle of each raw cone beam such that it is not much wider than the transversal cone angle of the associated cone beam proper. A group of detectors modules may be selected so that for a first of the radiation sources in the plurality of the radiation sources, and for at least a second of the radiation sources in the plurality of radiation sources, the first and the at least second group of radiation detector modules do not overlap on the detector. For example, for i = 1, ..., NJ2, the raw and proper cone beams of sources i and i + NJ2 may be provided so that they do not overlap on the detector.

FIG. 6 illustrates the cone beams proper of the sources 1 (marked by reference numeral 68), i (marked by reference numeral 67), and i + NJ2 n, for i close to NJA (marked by reference numeral 69).

Cone beams that start near one end of the source curve are a little asymmetric and have a slightly narrower transverse cone angle.

The intersection of the proper cone beams of all x-ray sources of the scanner forms the "volume of full projection;" which is the volume that is "seen" by all proper cone beams. With the acquisition geometry of the present invention the fully projected volume is large enough to contain a human heart and a good deal of its surroundings. In the example scanner, a sphere with a diameter of 20 cm is contained in the volume of full projection. Moreover, the fully projected volume is a subvolume of the convex hull of the source trajectory. This means that Tuy's completeness condition is satisfied in the volume of full projection. It is a clear advantage that Tuy's completeness condition is satisfied, since then this volume can be reconstructed without the "cone beam" artifacts that are typical of, and unavoidable with, circular and other planar source trajectories.

When scanning a patient' s volume of interest, a number of scan parameters need to be properly chosen. Such parameters include the tube voltage E and the tube current /. Also important is the sequence in which the sources are fired, the number N_0n of how often a source is switched on during a scan, and the duration T_0n of each on-time of each source. (The on- time of a source, as this term is used here, is not related with the pulsed working mode of carbon-nanotube-based x-ray sources). The parameters E, I, N_0n, and T_0n may all be set independently of the source number i. From the mentioned and other "primary" scan parameters, a number of "secondary" scan parameters can be derived. Among the secondary parameters are the total radiation time T_s = N_0n T_0n of each source; the total radiation time T_tot = N_ST_S of all sources; the power P = IE supplied to each source, which affects the heating of the source anode; and the time-current product T_totI, which is an indicator of the number of x- ray photons produced during the scan and which is thus related to both the radiation dose applied to the patient and to the signal-to-noise ratio of the reconstructed image. The total scan time r_scan, which is another secondary parameter, is related to the total radiation time, but may differ from the latter when there are idle times during the scan with no sources on, or when several sources are on simultaneously.

In conventional, single slice, single image, fan beam CT using a conventional rotating gantry with a single x-ray tube and a single detector row, a typical tube voltage is 120 kV, a typical tube current is 300 rriA, and a typical total radiation time is 500 ms, which is also the rotation time of the gantry. With these primary parameters, the applied power becomes 36 kW and the time-current product becomes 150 mAs. These parameters are also appropriate for the related "circular" version of cone beam CT which results from the above fan beam version when the single detector row is replaced by a multi-row detector with a large number of rows. Since the anode of the x-ray tube would quickly melt if it were static, the anode is made to rotate in order to dissipate the heat. It is a drawback of conventional CT gantries that they do not rotate quickly enough for a heart scan to be completed within 100 ms, and making them that fast would be very costly, if possible at all. With the cone beam CT scanner in accordance with embodiments of the present invention, it is desirable to achieve a similar time-current product, using a similar tube voltage. Again, the anodes of the x-ray sources would heat up very quickly, but now it is not possible to rotate the anodes to dissipate the heat. On the other hand, since the sources are individually addressable they may be switched on and off very rapidly, and in any order.

FIG. 7 illustrates a flow diagram of operating an imaging system in accordance with embodiments of the present invention.

Radiation is provided 70 from at least a first radiation source from the plurality of distributed radiation sources.

The radiation is detected 71 with at least a first radiation detector from the plurality of distributed radiation detectors, i.e. the associated detector group.

The at least first radiation source and the at least first radiation detector are operated 72 such that at least part of the radiation emitted from the at least first radiation source is detected by the at least first radiation detector.

In exemplary embodiments, the sources and detectors may be operated so that the radiation sources are turned on sequentially, and the group of radiation detectors associated with the radiation sources are read out sequentially. In the following, example embodiments of sequences or scan protocols are provided.

In an exemplary embodiment, the source sequence is set so that the sources fire one at a time in the sequential order i = 1, 2, ... N_s, with each source being on for JO_n seconds and a delay of Jl_di_e seconds between the firing of two neighbouring sources. A read sequence is set such that the detector is readout and reset during the delays. The total scan time of such a "sequential scan" is ⁷^_n ^{= N}* ^T™ ⁺ (^ ^{" 1}^ . In an alternative exemplary embodiment, interleaved sequences are used, with the sequential order i = 1, 1+ NJ2, 2, 2+ NJ2, ... (The number of sources is preferably even in this case). In the situation where the raw cone beams associated with sources i and i + NJ2 do not overlap on the detector, source and detector sequence may be set so that the reading of detector group i and the firing of source i+ NJ2 may overlap in time with the firing of source i and the reading of detector group i+ NJ2. The idle time Jl_di_e is thereby reduced. In the best case, there is no idle time at all, and the total scan time of such an "interleaved scan protocol" reduces from ^r— ^{= N} _* ^T _™ + (^S -ⁱfrdle. to ^can = NJ_m .

In yet another exemplary embodiment, pairs of sources (i, i+ NJ2) may be fired simultaneously, e.g. in the sequential order i = 1, 2, ..., NJ2. In this case, some idle time Ji_di_e should be introduced for reading the detector. Nevertheless, the total scan time of such a

"sequential double scan" reduces from ^{scan = N}s^T _™ ⁺ v*s ^{" 1}FIdIe to

T-_SCan = ^_0n /2+ 0v_s /2-i)r_ldle _

With the above or similar scan sequences, in order to arrive at a desired time current product, a relatively long on-time JO_n may be chosen. In the event that the anodes heat up too much during a specific on-time the originally planned single (sequential, interleaved, sequential double, or still other) scan sequence may be replaced by a faster scan sequence with a reduced on-time. A waiting time may be added between these faster scans. The measured results of the faster scans would be averaged to obtain the effect of the originally planned single scan. Splitting up a single "slow" scan into multiple faster scans in this way may reduce the temperature rise of the anodes by introducing an extra cool-down time for the anodes. The original time current product remains unchanged.

Other scan sequences than the ones described may be possible.

In many scan situations, the scanned part of the body must remain essentially static during the whole scan. For many parts of the body, when the patient simply lies still, the allowed maximum scan time is an ample plenty of seconds. The heart, however, rests only about 100 ms during a single heart beat. In such a situation, the sequence of the radiation sources being turned on may be correlated to an input signal, so that the sequence may be stopped upon receipt of a stop signal, and resumed upon a resume signal. Such an input signal may e.g. be provided by the patient's ECG signal, which may be used to start the scan precisely at the beginning of the resting period. If the time-current product that can be obtained within the available 100 ms even with the best possible scan sequence is deemed to be too small, one may repeat this scan sequence during the resting periods of several subsequent heart beats and average the measured results. The patient's ECG signal may be used to re- start the scan sequence during each heart beat. This scan mode requires the patient not only to lie still, but also to hold his or her breath for the required number of heart beats. In contrast to state-of-the-art cardiac helical CT, retrospective gating is not required, and the applied radiation dose is not increased compared to the dose that is required for the desired signal-to-noise ratio.

The acquired raw data may be pre-processed as usual in CT, e.g. by turning them into line integrals of the linear x-ray attenuation coefficient. Adverse effects of scattered x-rays may also be corrected or reduced using a suitable correction method. The reconstruction itself may be achieved using available reconstruction algorithms. If the reconstruction algorithm assumes a flat detector that is subdivided into rectangular array of detector elements, the measured cone beam projections may be re-sampled onto a virtual detector of the required type. Various pre-processing techniques and reconstruction algorithms are available to the skilled person.

The invention can be implemented in any suitable form including hardware, software, firmware or any combination of these. The invention or some features of the invention can be implemented as computer software running on one or more data processors and/or digital signal processors. The elements and components of an embodiment of the invention may be physically, functionally and logically implemented in any suitable way. Indeed, the functionality may be implemented in a single unit, in a plurality of units or as part of other functional units. As such, the invention may be implemented in a single unit, or may be physically and functionally distributed between different units and processors. Although the present invention has been described in connection with the specified embodiments, it is not intended to be limited to the specific form set forth herein. Rather, the scope of the present invention is limited only by the accompanying claims. In the claims, the term "comprising" does not exclude the presence of other elements or steps. Additionally, although individual features may be included in different claims, these may possibly be advantageously combined, and the inclusion in different claims does not imply that a combination of features is not feasible and/or advantageous. In addition, singular references do not exclude a plurality. Thus, references to "a", "an", "first", "second" etc. do not preclude a plurality. Furthermore, reference signs in the claims shall not be construed as limiting the scope.

Claims

1. A computed tomography imaging system comprising

- a plurality of distributed radiation sources (1, 20, 49) arranged to emit radiation beams from a plurality of emission locations, the plurality of emission locations being arranged substantially along a source trajectory, the radiation penetrating a volume of interest around an isocenter (3);

- a plurality of distributed radiation detector modules (4, 21, 54, 55), the plurality of radiation detector modules being arranged substantially along a detector trajectory;

wherein the source and detector trajectories are such that a straight line (8) that starts at a point on the source trajectory and passes through the isocenter intersects the detector trajectory; and wherein the radiation beam that is emitted from an emission location on the source trajectory is not significantly blocked by a portion of a radiation detector module between the emission location and the volume of interest (5).

2. The imaging system according to claim 1, wherein the source trajectory and the detector trajectory are defined by the general form of: d (λ) = sr_t (cos α(λ) cos β(λ), sin β(λ),- sin α(λ) cos β(λ)) where, r_u defines a distance from the isocenter of the volume of interest, s is either 1 or -1, and a(λ), β(λ) are functions (30, 31) of

3. The imaging system according to claim 1, wherein the plurality of distributed radiation sources (20) ranges from 500 to 1000 radiation sources.

4. The imaging system according to claim 1, wherein the radiation sources are carbon- nanotube-based x-ray sources (49).

5. The imaging system according to claim 1, wherein a sensitive area of the odd and even numbered detector modules are arranged alternatively with a first and second distance (52, 53) from the isocenter of the volume of interest.

6. The imaging system according to claim 5, wherein a cross-section of the detector modules is rectangular (56) or T-shaped (57).

7. The imaging system according to claim 1, wherein the plurality of distributed radiation sources and the plurality of distributed detector modules are individually addressable.

8. A method of operating a cone beam computed tomography imaging system, the method comprising the steps:

- providing radiation (70) from at least a first radiation source from a plurality of distributed radiation sources arranged to emit radiation beams from a plurality of emission locations, the plurality of emission locations being arranged substantially along a source trajectory, the radiation penetrating a volume of interest around an isocenter;

- detecting radiation (71) with at least a first radiation detector module, the at least first radiation detector module being part of a plurality of distributed radiation detector modules, the plurality of radiation detector modules being arranged substantially along a detector trajectory;

the source and detector trajectories being such that a straight line that starts at a point on the source trajectory and passes through the isocenter intersects the detector trajectory; and wherein the radiation beam that is emitted from an emission location on the source trajectory is not significantly blocked by a portion of a radiation detector module between the emission location and the volume of interest; and

where the at least first radiation source and the at least first radiation detector module are operated (72) such that radiation emitted from the at least first radiation source passes through the volume of interest and is detected by the at least first radiation detector module.

9. The method according to claim 8, wherein each distributed radiation source and each distributed detector module are individually addressable.

10. The method according to claim 8, wherein for each radiation source of the plurality of radiation sources, a group of radiation detector modules are associated with the radiation source so as to detect at least a part of the emitted radiation beam, and wherein the group of detector modules are selected so that for a first of the radiation sources in the plurality of the radiation sources, and for at least a second of the radiation sources in the plurality of radiation sources, the first and the at least second group of radiation detector modules do not overlap.

11. The method according to claim 8, wherein the radiation sources are turned on sequentially, and wherein the group of radiation detector modules associated with the radiation sources are read out sequentially.

12. The method according to claim 11, wherein the sequence of the radiation sources being turned on, is correlated to an input signal, so that the sequence may be stopped upon receipt of a stop signal, and resumed upon a resume signal.

13. A computer program product having a set of instructions, when in use on a computer, to cause the computer to perform the steps of claim 8.