Patent Application
20160073993
The present disclosure relates generally to medical imaging and, more particularly, to systems and methods for tracking motion during medical imaging procedures using motion tracking coils.
Positrons are positively charged electrons which are emitted by radionuclides that have been prepared using a cyclotron or other device. These are employed as radioactive tracers called “radiopharmaceuticals” by incorporating them into substances, such as glucose or carbon dioxide. The radiopharmaceuticals are administered to a patient and become involved in biochemical or physiological processes such as blood flow; fatty acid and glucose metabolism; and protein synthesis.
As the radionuclides decay, they emit positrons. The positrons travel a very short distance before they encounter an electron, and when this occurs, they are annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features which are pertinent to PET scanners—each gamma ray has an energy of 511 keV and the two gamma rays are directed in nearly opposite directions. An image indicative of the tissue concentration of the positron emitting radionuclide is created by determining the number of such annihilation events at each location within the field of view.
A conventional positron emission tomography (PET) imaging system includes one or more rings of detectors which encircle the patient and which convert the energy of each 511 keV photon into a flash of light that is sensed by a photomultiplier tube (PMT). Coincidence detection circuits connect to the detectors and record only those photons which are detected simultaneously by two detectors located on opposite sides of the patient. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors. Within a few minutes hundreds of million of events are recorded to indicate the number of annihilations along lines joining pairs of detectors in the ring. These numbers are employed to reconstruct an image using well known computed tomography techniques.
Many minutes are typically required to accumulate a sufficient number of counts in a PET imaging system in order to reconstruct an image having sufficient SNR to be of clinical value. During that time period the subject of the examination is prone to move at least one or more times. As a result, the image that is reconstructed is often blurred.
For example, PET scans are an important part in the diagnosis, prognostication, and monitoring of dementia, which represents a patient group that is particularly susceptible to motion. That is, artifacts from head motion are the major challenge facing brain PET. This is particularly true in the elderly, with increased significance in patients with dementia or movement disorders. Also, voluntary patient body (other than brain) motion is also common in clinical body PET imaging (studies shown 29% prevalence in dynamic myocardial perfusion PET), which degrades the quality of the PET images and affect the diagnostic value of the PET exam, possibly leading to misdiagnosis.
Many approaches have been explored in the effort to correct motion artifacts. Depending on whether the motion is estimated from the acquired PET data or by other instrumentation, the approaches can be divided into two groups: auto-correction and assisted-correction. For the auto-correction techniques, the measured PET data are divided into temporal frames, and the motion is then estimated between temporal frames from the PET data. The estimated motion field can then be used to transform the reconstructed images (Friston et al., 1995; Tellmann et al., 2006; Woods et al., 1992) or the sinograms (Hutton et al., 2002; Kyme et al., 2003) of each temporal frame to a reference frame. The accuracy of motion estimation using this approach is limited by the noise of PET images, which increases as the data set is divided into temporal frames for a dynamic image sequence. Moreover, the fact that the motion estimation relies on the generation of images or sinograms limits its temporal resolution. Thus, such methods are not suitable when the activity distribution is fast changing or the object is fast moving. The reconstruction algorithms of the assisted-correction approaches are similar to auto-correction techniques except that the motion information is instead measured using an instrument other than the PET camera, such as video/infrared cameras (Bloomfield et al., 2003; Goldstein et al., 1997; Picard and Thompson, 1997), and approaches with structured light (Olesen et al., 2012, 2013).
Similar approaches have also been applied to motion correction in MRI (Schulz et al., 2012; Zaitsev et al., 2006). One advantage of the optical motion tracking approaches is that they are independent of the MRI acquisition, so that no changes to the MR pulse sequence are required, in contrast to MR navigator-based methods. Another advantage is that the optical methods, in principle, are capable of achieving high frame rate. Some of these approaches monitor the motions of reflectors attached to the subject's head and some observe a portion of the subject's face. In every case, these methods require an unobstructed view from the cameras to the reflectors or the subject's face. This is challenging for combined PET-MR systems performing head studies because the view from outside of the scanner is blocked by the MR head coil, especially for modern head coils with a large number of channels. There are RF contamination and MR compatibility issues associated with installing cameras inside of the scanner. Moreover, these optical systems require complicated calibrations.
Conventional MR navigator methods (Ehman and Felmlee, 1989; Wang et al., 1996) can be used to track motion with temporal resolution less than 20 ms. However, such methods cannot be used to track head rotation. Catana et al. (2011) used the cloverleaf navigator method (van der Kouwe et al., 2006) to track head motion for PET motion compensation. Although this method can track both translation and rotation, its motion tracking accuracy suffers from off-resonance effects, gradient instabilities, as well as signal contamination from non-rigid motion of the neck. Moreover, such cloverleaf-navigator methods require approximately 20 s of motionless data to calibrate. Petibon et al. (2013) used image-based MR motion tracking for non-rigid motion compensation in cardiac PET. However, it generally lacks temporal resolution because of the long scanning time needed for acquiring the entire image volume.
Therefore, there is a need for systems and methods to control or offset motion, particularly, in PET imaging.
The present disclosure provides system and methods that overcome the aforementioned and other drawbacks. Particularly, a plurality of coils, tracked by magnetic resonance imaging (MRI), are used to determine head or body motion in real-time during a PET acquisition and incorporate the measured motion in the PET imaging reconstruction process or apply the measured motion to the corresponding PET image reconstructed at each motion frame. Each tracking coil can be wound around a sealed doped water sample to receive only the signal from the sample. A projection of the sample contains a single sharp peak indicating the location of the tracking coil along the field gradient. Thus, three orthogonal projections can be used to yield the tracking coil position in 3D space, which can be performed at high speed. The temporal resolution for tracking coil tracking can be better than 15 ms. The motion measured by MR tracking coils can then be used for motion correction in the PET image reconstruction process.
In accordance with one aspect of the disclosure, a medical imaging system is disclosed that includes a positron emission tomography (PET) system for acquiring a series of medical images of a subject, the PET system. The PET system includes a plurality of detectors arranged about a bore configured to receive the subject and to acquire gamma rays emitted from the subject as a result of a radiotracer administered to the subject and configured to communicate PET signals corresponding to acquired gamma rays. The medical imaging system also includes a magnetic resonance imaging (MRI) system for acquiring a series of medical images of the subject from within the bore. The MRI system includes a magnet system configured to generate a polarizing magnetic field about at least a portion of the subject arranged in the bore and a plurality of gradient coils configured to apply a gradient field to the polarizing magnetic field. The MRI system also includes a radio frequency (RF) system having a plurality of motion tracking coils positioned, the RF system configured to acquire MR data from the motion tracking coils, wherein each motion tracking coil includes a conductor extending through a plurality of loops to form a coil. The MRI system further includes an MRI computer system programmed to control the RF system to receive MR data from the RF system. The medical imaging system also includes a data processing system configured to receive the MR data and determine motion of the subject from the MR data by determining a position of the motion tracking coils over time and receive the PET signals and correct the PET signals using the motion of the subject determined from the MR data to generate corrected PET images of the subject.
In accordance with another aspect of the disclosure, a method is disclosed for generating a positron emission tomography (PET) image corrected for subject motion in a combination PET and magnetic resonance imaging (MRI) system. The method includes preparing a subject to be imaged in a combined PET and MR imaging process by administering a radionuclide to the subject and arranging a plurality of motion tracking coils about the subject, wherein each of the plurality of motion tracking coil includes a conductor extending through a plurality of loops to form a coil configured to receive signals from a sample material. The method also includes acquiring, with a PET imaging system, sinogram data that counts a number of coincidence events at a plurality of lines of response (LOR) and periodically acquiring, with the MRI system, over a time period that extends concurrently with acquiring the sinogram data, a plurality of MR signals generated using the plurality of motion tracking coils arranged about the subject. The method also includes calculating, from the MR data, motion data and correcting the sinogram data using the motion data to generate corrected PET images of the subject.
The foregoing and other aspects and advantages of the disclosure will appear from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown by way of illustration a preferred embodiment of the invention. Such embodiment does not necessarily represent the full scope of the invention, however, and reference is made therefore to the claims and herein for interpreting the scope of the invention.
Referring to
Referring particularly to
The event locator circuits 227 form part of a data acquisition processor 230, which periodically samples the signals produced by the acquisition circuits 225. The processor 230 has an acquisition CPU 229 which controls communications on local area network 218 and a backplane bus 231. The event locator circuits 227 assemble the information regarding each valid event into a set of digital numbers that indicate precisely when the event took place and the position of the scintillator crystal which detected the event. This event data packet is conveyed to a coincidence detector 232 which is also part of the data acquisition processor 230.
The coincidence detector 232 accepts the event data packets from the event locators 227 and determines if any two of them are in coincidence. Coincidence is determined by a number of factors. First, the time markers in each event data packet must be within a preset time of each other, and second, the locations indicated by the two event data packets must lie on a straight line. Events that cannot be paired are discarded, but coincident event pairs are located and recorded as a coincidence data packet. As will be described, the coincidence data packets can be corrected for motion of the subject during the acquisition using information received from the MRI system 300 of
The corrected coincidence data packets are conveyed through a link 233 to a sorter 234 where they are used to form a sinogram. This corrective process is repeated each time corrective values are received from the MRI system. The correction is made on those coincidence data packets that have accumulated since the receipt of the previous corrective values.
The sorter 234 forms part of an image reconstruction processor 240. The sorter 234 counts all events occurring along each projection ray (R, θ) and organizes them into a two dimensional sinogram array 248 which is stored in a memory module 243. In other words, a count at sinogram location (R, θ) is increased each time a corrected coincidence data packet at that projection ray is received. Due to the corrections made to the coincidence events, the sinogram that is formed during the scan depicts the subject being examined in the reference position despite subject motion that occurs during the scan. The image reconstruction processor 240 also includes an image CPU 242 that controls a backplane bus 241 and links it to the local area network 218. An array processor 245 also connects to the backplane 241 and it reconstructs an image from the sinogram array 248. The resulting image array 246 is stored in memory module 243 and is output by the image CPU 242 to the operator work station 215.
The operator work station 215 includes a CPU 250, a display 251 and a keyboard 252. The CPU 250 connects to the network 218 and it scans the keyboard 252 for input information. Through the keyboard 252 and associated control panel switches, the operator can control the calibration of the PET scanner and its configuration. Similarly, the operator can control the display of the resulting image on the display 251 and perform image enhancement functions using programs executed by the work station CPU 250.
Referring to
The pulse sequence server 310 functions in response to instructions downloaded from the operator workstation 302 to operate a gradient system 318 and a radiofrequency (RF) system 320. Gradient waveforms necessary to perform the prescribed scan are produced and applied to the gradient system 318, which excites gradient coils in an assembly 322 to produce the magnetic field gradients Gx, Gy, and Gz used for position encoding magnetic resonance signals. The gradient coil assembly 322 forms part of a magnet assembly 324 that includes a polarizing magnet 326 and optionally a whole-body RF coil 328.
RF waveforms are applied by the RF system 320 to the RF coil 328, or a separate local coil, such as the tracking coil system 400, as illustrated in
The RF system 320 also includes one or more RF receiver channels. Each RF receiver channel includes an RF preamplifier that amplifies the magnetic resonance signal received by the coil 328 or by the local coil or optionally by a coil in the coil array 400 to which it is connected, and a detector that detects and digitizes the I and Q quadrature components of the received magnetic resonance signal. The magnitude of the received magnetic resonance signal may, therefore, be determined at any sampled point by the square root of the sum of the squares of the I and Q components M=√{square root over (I2+Q2)} and the phase of the received magnetic resonance signal may also be determined according to the following relationship
The pulse sequence server 310 also optionally receives patient data from a physiological acquisition controller 330. By way of example, the physiological acquisition controller 330 may receive signals from a number of different sensors connected to the patient, such as electrocardiograph (ECG) signals from electrodes, or respiratory signals from a respiratory bellows or other respiratory monitoring device. Such signals are typically used by the pulse sequence server 310 to synchronize, or “gate,” the performance of the scan with the subject's heart beat or respiration.
The pulse sequence server 310 also connects to a scan room interface circuit 332 that receives signals from various sensors associated with the condition of the patient and the magnet system. It is also through the scan room interface circuit 332 that a patient positioning system 334 receives commands to move the patient to desired positions during the scan.
The digitized magnetic resonance signal samples produced by the RF system 320 are received by the data acquisition server 312. The data acquisition server 312 operates in response to instructions downloaded from the operator workstation 302 to receive the real-time magnetic resonance data and provide buffer storage, such that no data is lost by data overrun. In some scans, the data acquisition server 312 does little more than pass the acquired magnetic resonance data to the data processor server 314. However, in scans that require information derived from acquired magnetic resonance data to control the further performance of the scan, the data acquisition server 312 is programmed to produce such information and convey it to the pulse sequence server 310. For example, during prescans, magnetic resonance data is acquired and used to calibrate the pulse sequence performed by the pulse sequence server 310. As another example, navigator signals may be acquired and used to adjust the operating parameters of the RF system 320 or the gradient system 318, or to control the view order in which k-space is sampled. In still another example, the data acquisition server 312 may also be employed to process magnetic resonance signals used to determine patient motion, as will be described, and communicate such to the PET system 200 described with respect to
The data processing server 314 receives magnetic resonance data from the data acquisition server 312 and processes it in accordance with instructions downloaded from the operator workstation 302. Such processing may, for example, include one or more of the following: reconstructing two-dimensional or three-dimensional images by performing a Fourier transformation of raw k-space data; performing other image reconstruction algorithms, such as iterative or backprojection reconstruction algorithms; applying filters to raw k-space data or to reconstructed images; generating functional magnetic resonance images; calculating motion or flow images; and so on.
Images reconstructed by the data processing server 314 are conveyed back to the operator workstation 302 where they are stored. Real-time images are stored in a data base memory cache (not shown in
The MRI system 300 may also include one or more networked workstations 342. By way of example, a networked workstation 342 may include a display 344; one or more input devices 346, such as a keyboard and mouse; and a processor 348. The networked workstation 342 may be located within the same facility as the operator workstation 302, or in a different facility, such as a different healthcare institution or clinic.
The networked workstation 342, whether within the same facility or in a different facility as the operator workstation 302, may gain remote access to the data processing server 314 or data store server 316 via the communication system 340. Accordingly, multiple networked workstations 342 may have access to the data processing server 314 and the data store server 316. In this manner, magnetic resonance data, reconstructed images, or other data may exchanged between the data processing server 314 or the data store server 316 and the networked workstations 342, such that the data or images may be remotely processed by a networked workstation 342. This data may be exchanged in any suitable format, such as in accordance with the transmission control protocol (TCP), the internet protocol (IP), or other known or suitable protocols.
Referring to
The water sample 408 inside the housing 406 may be doped to yield short T1 and/or T2 relaxation times. To this end, in the non-limiting example provided above, the sphere can be filled with a degassed solution of deionized water doped with 1.25 g/L NiSO4.6H2O and 5 g/L 180 NaCl. The sphere can be immersed in liquid nitrogen while the air is pumped out, and the neck was flame-sealed under vacuum. The tracking coil 404 may wrap around the housing 406 and be connected to one or more capacitors 410 to resonate the coil at the Lamor frequency of the MR system, such as the MR system 300 of
As illustrated in
If communicating wirelessly using magnetic inductive coupling no impedance matching circuitry is needed. The capacitor 410 may be adjustable to allow the coil system 400 to be fine-tuned with the water sample 402 in place by observing the resonance frequency while the coil system 400 is probed with a coupling loop connected, for example, to a Hewlett-Packard 8753C vector network analyzer displaying reflected power. Alternatively, the capacitor 410 may be a fixed value capacitor, and the coil circuit resonance frequency fine-tuned by adjusting the spacing of the coil windings. In another alternative, the coil circuit resonance frequency may be fine-tuned by adjusting the capacitor and winding spacing in combination.
The typical tracking coil quality factor Q (a dimensionless measure of the sharpness of coil tuning) was tested and, in the above-described non-limiting example, was on the order of 120 with the water sample inserted. The coil system 400 was enclosed in polypropylene tubes to prevent contact with the subject. Additional tracking coils were built similarly with finer gauge, e.g., 26 AWG, insulated magnet wire using a larger number of turns in a close-wound solenoid. The tracking coils using finer wire tended to exhibit higher Q but less stability in tuning, presumably because the finer wire was less rigid, this permitted inductance changes during handling. None of the tracking coils in this study included back-to-back switching diodes to limit the RF excitation applied to the water sample, because the diodes tended to lower the Q as well as the detection sensitivity. In some cases propylene tubes (syringe barrels) were used to support the tracking coil from the inside and provide a “handle” so that the coils could be handled or fixed to the object being imaged.
In operation, the coil system 400 can work by inductive coupling with the coil 328 of the RF system 320 (and with the separate receiver coil if used). If the coil system 400 is not wired to the MRI system 300, but rather communicates wirelessly by inductive coupling, the RF magnetic field of the transmit pulse excites an RF current in the tracking coil 404 of the coil system 400, which in turn excites the spins in the water sample 402. Although the mutual inductance between the coil 328 of the RF system 320 and the tracking coil system 400 is quite small, the RF field (B1) produced in the water sample 402 is quite large because of the large filling factor and the high quality factor (Q-factor) of the tracking coil 404, such that an RF flip angle of only a fraction of a degree (measured with respect to the body coil for the imaged subject, not the tracking sample) excites an intense signal from the coil system 400. Thus, mutual inductance between the coil system 400 and the coil 328 of the RF system 320 couples the MR signal of the water sample 402 into the receiver, yielding a readily detectable signal.
The location of the coil system 400 can be readily tracked using a variety of MR pulse sequences. However, as will be described, one particular pulse sequence includes 3-6 short-duration RF excitations followed by a gradient-echo readout that yields the projections of the markers onto certain axes. The projections (often X, Y and Z projections) are then combined to generate the 3D location of the marker in the MRI system.
Referring to
Although intuitively, it might be thought that the trajectory to acquire the volume projections needs to go through the center of the k-space, the application of the spoiler gradient 510 effectively translates the k-space position in the direction perpendicular to the tracking readout direction.
In the tracking pulse sequence module 500 illustrated in
As illustrated in
The pre-positioning 504′ and readout gradients 506′ are also used as spoiler gradients. This is similar to the tracking pulse sequence 500 of
Referring to
Thus, the tracking sequence module 500 of
Thus, the tracking sequence module 500′ of
Hence, as described, present disclosure provides a pulse sequence that does not rely on “navigators” or a “navigator sequence” that rapidly acquires a crude low spatial resolution version of the MR image, for example a simple 1D projection of the subject, and or methods that use such navigator data to either correct the incoming raw data as it is acquired, or to adjust the scanner operation (such as shifting the image plane), or else is saved for motion correction in post processing. Rather, as described, MR data is acquired from a set of tracking coils. To this end, the above-described methods can be implemented even if MR imaging (including navigator imaging) of the subject is not being conducted, and even if the MR apparatus (transmit and receive coils) that are required for MR imaging are absent from the MRI scanner. Further, the systems and method of the present disclosure do not require or use of image data from the subject and, as described, can optionally use data acquired from a sample that is separate from the subject.
To demonstrate the utility of the proposed efficient trajectory, a phantom study was performed in the context of MR active marker based motion correction for PET imaging in simultaneous MR-PET acquisition. The acquisitions were performed on a simultaneous MR-PET scanner (Siemens Biograph mMR, Siemens Healthcare, Erlangen, Germany).
Referring to
More particularly, motion corrected PET image reconstruction will now be described. Several methods have been proposed for PET motion correction. One method is to divide detected events into multiple static motion phases using a specified threshold and perform individual reconstruction in each phase. This is then followed by registering each frame to a reference motion phase and summing all the registered images. However, a high threshold can cause the motion within the phase to be ignored. On the other hand, a low threshold can lead to low statistical phase to be reconstructed. Lack of an adequate number of acquired events in the individual phases can, in turn, adversely affect the quality of the final reconstructed images. Moreover, an increased number motion phases can lead to increased computation time. Another approach is to perform deconvolution on the motion-blurred reconstructed images. However, the deconvolution amplifies the noise in the PET data. When the motion is significant, it also requires spatially variant deconvolution filters, which increase computational costs and introduce other artifacts. Another method is to model the motion in the PET system matrix in PET reconstruction. This method is usually applied to non-rigid motion. Although such a method can also be applied to rigid motion, it would converge slowly. For rigid motion, such as head motion, the most accurate approach is to correct the individual lines of response.
Most regions of interest studied in PET, such as the head, undergoes rigid body motion. Thus, the motion of the object can be uniquely determined by the locations of multiple non-collinear active markers, such as using the above-described tracking coil systems illustrated in
where xjk is the activity in voxel i (i=1,2, . . . ,J) at the kth iteration,
In particular,
Testing results show that the efficient trajectory for active marker tracking of the present disclosure provides fast and accurate marker location measurements. Using the efficient trajectory, the duration of the active marker tracking sequence module has been shortened to 5.5 ms. Sub-5ms duration of the tracking module can be achieved with slightly varied parameters: pixel size=1.52 mm, 256 samples per line, 2 times readout oversampling, bandwidth=1502 Hz/pixel, allowing an update rate higher than 200 FPS; whereas, to our knowledge, the reported shortest duration of the conventional tracking sequence module in the literature is 12.9 ms10.
In one set of experiments, the body RF coil was used to wirelessly receive the signal emitted by the active markers. In practice, bird-cage or multichannel head coils can be used to boost the SNR of the signals from the markers. For multichannel coils, combination of data from the channels can be performed by a simple sum-of-squares.
Certain errors may occur during position tracking with coils or microcoils. For example, due to the presence of copper wires and glass in the vicinity of the doped water samples in the active markers, a local magnetic susceptibility induced field shift is expected. The water sample might not be perfectly spherical because the water might enter the neck of the cell, or a bubble caused by insufficient filling may be present; this effect will create a small susceptibility shift, as well as possible broadening and asymmetry of the profile. Both of these sources of shift will exhibit a dependence on orientation of the marker.
The present disclosure studied how the susceptibility affects the accuracy of the location measurements. This effect can be measured by repeating the tracking module with opposite gradient polarity. This study was performed on one wireless active marker with the same acquisition parameters as used in the motion correction study except the readout bandwidths were varied (130, 201, 300, 592, 814 and 1149 Hz/pixel). The readout bandwidth is proportional to the magnitude of the readout gradient. A stronger readout gradient yields a more accurate position because it dominates small field shifts caused by other effects such as magnetic susceptibility, but reduces the SNR.
The difference between the centroids of the X, Y and Z projections read out with gradients of opposite polarity were calculated for each readout bandwidth and the results are shown in
When lower bandwidth is used, the efficiency improvement of the proposed trajectory will be less. The efficiency improvement is largest for high bandwidth since the proposed trajectory improves the efficiency by reducing time spent on magnetization preparation. The magnetization preparation time will not change much when the bandwidth is reduced, but the ADC duration will be longer. However, in practice, the highest achievable bandwidth is almost always desired because the sampling time needed for the tracking module increases with the decrease of bandwidth.
Although it is intuitive to assume perfect spatial alignment between PET and MR, the spatial accuracy of MR is affected by gradient non-linearity induced spatial distortion especially at the locations far from the scanner isocenter. When not corrected, the spatial distortion can lead to misalignment between the attenuation map and emission image, and error in motion tracking. The gradient non-linearity was measured when the scanner was installed and the parameters are stored on the scanner, allowing one to account for it. The spatial distortion correction can be performed off-line (Jovicich et al., 2006). The spatial distortion correction was not performed in these studies because the markers were close to the isocenter of the scanner, where the distortion caused by gradient non-linearity was small. Also, due to the small displacements of the markers, gradient non-linearity had only second-order effects on the accuracy of our measurements.
The MR tracking sequence module shown in
In some tests, the traces from the tracking coils do not overlap. In practice, the traces may cross each other if the markers are not placed far apart from each other. When that happens, the correspondence between the markers and the traces can still be determined by utilizing the constraint that the distances between the active markers remain the same under rigid motion. One measurement generates three locations on each of the X, Y and Z projections. In total, there are 33=27 possible 3D locations for the three markers. Of these, there are 36 possible combinations forming a triangle, as they must. The triangle with the edge lengths that match with the initial placement will be selected. Moreover, if the locations were frequently measured (as was done in this work), the constraint that the displacement of each marker is small between two measurements can also be used for the correspondence determination.
The motion tracking and correction with the proposed efficient trajectory was demonstrated in retrospective motion corrected PET image reconstruction. However, the techniques described herein can also be used in prospective motion correction in MR imaging or catheter tracking. The improved acquisition efficiency can reduce the overhead for motion tracking or improve the temporal resolution.
Thus, a system and method for efficient active marker tracking is provided for correcting PET data using MR data in a combined or simultaneous MR-PET scanner. The tracking sequence module based on the proposed trajectory can obtain the locations of multiple active markers within 5 ms and is significantly more efficient than conventional tracking sequence modules. The improved efficiency of the tracking sequence module leads to less overhead for motion tracking and potentially higher frame rates for rapid motion.
It should be apparent that many variations from the above embodiments are possible without departing from the invention. For example, the MRI system and PET scanner may be more fully integrated with control and processing components being shared by both systems. As another example, the material detected by the tracking coils may be contained with the coils or may be adjacent to the exterior of the coils. As another example, the tracking coils may be wired to the MRI scanner or may operate wirelessly via inductive coupling to the body RF coil or another coil in the scanner. As still another example, wired tracking coils may be used in transmit/receive mode, whereby the RF excitation is provided through the wired connection rather than via the body RF coil. As yet another example, the tracking coils can have winding configurations other than solenoids.
This application is based on, claims priority to, and incorporates herein by reference U.S. Provisional Application Ser. No. 61/817,750, filed Apr. 30, 2013, and entitled, “Motion Tracking Micro-Coils.”
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US14/36132 | 4/30/2014 | WO | 00 |
| Number | Date | Country | |
|---|---|---|---|
| 61817750 | Apr 2013 | US |
