Abstract:
A technique for acquiring magnetic field maps simultaneously with the images they affect allows improved correction of shimming and/or geometric distortions in the image and allows imaging techniques where subject motion is inevitable or required.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
   This application claims the benefit of U.S. Provisional Application 60/469,958 filed May 13, 2003 and hereby incorporated by reference. 

   STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT 
   This invention was made with United States government support awarded by NIH grants EB002014 and EB000215. The United States Government has certain rights in this invention. 

   BACKGROUND OF THE INVENTION 
   The present invention relates to magnetic resonance imaging (MRI) and functional magnetic resonance imaging (fMRI) and in particular to a pulse sequence and methodology for acquiring image data contemporaneously with information about variations in magnetic field homogeneity. 
   Magnetic resonance imaging is used to generate medical diagnostic images by measuring faint radio frequency (RF) signals (magnetic resonance) emitted by atomic nuclei in tissue (for example, water protons) after radio frequency stimulation of the tissue in the presence of a strong magnetic field. 
   The location of the precessing protons is made possible by the application of orthogonal magnetic gradient fields which serve to “encode” the spins according to frequency, phase, and/or slice. The combination of the radio frequency stimulation and the applied gradient fields is termed a pulse sequence. 
   The acquired signal from the spins (termed a nuclear magnetic resonance (NMR) signal) provides data in “k-space”, a mathematical construction in the frequency domain. A two-dimensional Fourier transform of the k-space data produces the actual image. It will be understood, therefore, that the k-space data does not represent the image itself, but represents the spectral components of the image with the center of k-space representing low frequency spatial components of the image, and the outer portions of k-space representing the high frequency spatial components of the image. 
   The impressing of spatial location information onto the spins of the NMR signal by the applied magnetic gradients makes it extremely important that all applied magnetic fields (including the polarizing magnetic field B 0  and the gradient magnetic fields G x , G y , and G z ) be well characterized. For this reason, and particularly for the B 0  field, it is well known to incorporate shimming coils into the design of a magnetic resonance imaging machine which serve to correct for inhomogeneities in the B 0  field through the application of one or more superimposed shimming fields. 
   A number of techniques are known by which to measure inhomogeneities of the magnetic field and thus to calculate the currents needed for the shimming coils. For example, special pulse sequences detecting phase differences in the MRI measurements of a homogenous phantom, for example, a tank of water, may be used to deduce variations in the magnetic field of the MRI system. 
   Shimming of the MRI system may be accomplished with great precision, however, the magnetic homogeneity is upset almost immediately upon insertion of a human subject whose tissue distorts the field. In order to address this problem, it is known to create a magnetic field map once a subject is in position in the MRI machine to compensate for this distortion. Such compensation is particularly important for echo planar imaging (EPI) and spiral imaging where the precessing nuclei have a long period of time in which to be influenced by the magnetic field, and thus to accumulate errors caused by inhomogeneity. 
   BRIEF SUMMARY OF THE INVENTION 
   The present inventors have recognized that while magnetic field maps acquired while the subject is in position in the MRI machine can improve static  shimming of the MRI machine, these magnetic field maps have two important shortcomings. The first is that the different pulse sequences used to acquire the image and used to acquire the magnetic field map produce different geometric distortions and thus are hard to register. The second is that subject movement between acquisition of the image and the magnetic field map can render the magnetic field map inaccurate with respect to distortions in the image. 
   Accordingly, the present inventors have developed a technique in which field homogeneity information is collected essentially at the same time as image data and using the same pulse sequence. As a result, changes in magnetic fields accompanying subject movement may be easily matched to the image and the image readily corrected. 
   Specifically then, the present invention provides an MRI imaging pulse sequence where a magnetic field map and image data are collected contemporaneously. For example, both the magnetic field map and the image may be acquired after a single, common RF pulse. 
   It is thus one object of at least one embodiment of the invention to provide image and magnetic field map data that are for all practical purposes coincident in time. 
   The magnetic field map may be acquired during the low order k-space line acquisitions of the image data. 
   Thus it is another object of at least one embodiment of the invention to make use of the same gradient fields to acquire the image. 
   Only low order k-space lines need be obtained for the magnetic field map. 
   It is thus another object of at least one embodiment of the invention to reduce acquisition time by limiting the magnetic field map to low spatial frequency variations. More generally, the invention allows the amount of redundant data collected for the field map to be traded for acquisition time. If a short acquisition time is required, only a few lines of redundant data can be collected. If longer times are permitted, then a larger number of redundant lines of data can be collected. 
   The pulse sequence may be a partial k-space acquisition and the full k-space data set may be mathematically extrapolated. 
   Thus it is another object of at least one embodiment of the invention to provide another method of allowing the amount of data collected to be traded for acquisition time. 
   The magnetic field map may be used to correct for susceptibility distortions in the associated image. 
   Thus it is another object of at least one embodiment of the invention to provide concurrent correction of image data for susceptibility-induced distortions. Although susceptibility induced distortion arises during patient motion, susceptibility distortion cannot be corrected by standard motion correction techniques that rely on simple translations and rotations of the acquired images. The invention therefore provides a technique useful to augment motion correction for acquisitions where subject motion is expected or intended, such as fMRI studies where the subject speaks, and where even small amounts of distortion can prevent proper image registration. 
   Multiple sets of image data and magnetic field map data may be acquired over time and an earlier magnetic field map may be used to provide control of the shimming coils of the MRI system to correct later images. The later magnetic field map may be then used to further correct the image for susceptibility effects. 
   Thus it is another object of at least one embodiment of the invention to allow for real time physical shimming of the magnetic field with and without later susceptibility corrections. 
   The magnetic field map may be used to discard particular images in a time series having excessive susceptibility error. 
   Thus it is a further object of the invention to use the magnetic field map to cull bad image data from sets of images. 
   These particular objects and advantages may apply to only some embodiments falling within the claims and thus do not define the scope of the invention. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
       FIG. 1  is a simplified schematic representation of a magnetic resonance imaging apparatus showing a coil driving and acquisition circuitry connected to a controller controlling the MRI system and communicating with a user interface; 
       FIG. 2  is a pulse sequence suitable for use with the present invention to collect both image and magnetic field map data; 
       FIG. 3  is a representation of k-space showing a “moving racetrack” trajectory in k-space providing two traversals of the same line of k-space for magnetic field and image data collection; 
       FIG. 4  is a figure similar to that of  FIG. 3  showing an alternative trajectory in k-space less delay between the two traversals of the same lines of k-space; 
       FIG. 5  is a schematic representation of the processing of the data acquired in  FIGS. 3 and 4 ; 
       FIG. 6  is a flow chart showing the steps of a program executed by the controller of  FIG. 1  in the present invention; and 
       FIG. 7  is a schematic representation of k-space showing a partial k-space acquisition. 
   

   DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
   Referring now to  FIG. 1 , magnetic resonance imaging machine  10  includes a magnet  12  providing for a substantially uniform B 0  field  14  within a bore  16  that may hold a subject. The magnet  12  supports a radio frequency coil (not shown) that may provide a radio frequency excitation to protons of the subject (not shown) within the bore  16 . The radio frequency coil communicates with an RF pulse subsystem  18  producing the necessary electrical waveform as is understood in the art. 
   The magnet  12  also supports three axes of gradient coils (not shown) of a type known in the art which communicate with a corresponding gradient subsystem  20  providing electrical power to the gradient coils to produce gradient coil functions G x , G y , and G z  over time. Also supported by the magnet  12  are multiple shimming coils (not shown) which communicate with shimming subsystem  22  to produce shimming magnetic fields within the bore  16 . 
   Finally, an NMR acquisition subsystem  24  connects to RF reception coils positioned within the magnet  12 . 
   Each of the subsystems  18  through  24  communicates with a central controller  26  which generates pulse sequences comprised of RF pulses from RF subsystem  18  and gradient pulses from gradient coil subsystem  20 , and which receives NMR signals through NMR acquisition subsystem  24 . The controller  26  operates according to one or more stored control programs  28  that may define the pulse sequences and which may operate on data  30  collected from the NMR acquisition subsystem  24  to produce images that may be displayed on console  32 . The console  32  also allows the input of data including programs and commands to the controller  26 . 
   Referring now to  FIG. 2 , in an echo planar gradient recalled imaging sequence suitable for use with the present invention, an RF excitation pulse  34  is generated at the beginning of the pulse sequence using the RF subsystem  18 . During the excitation, a G z  slice encoding gradient  36  may be generated along the z-axis (aligned with B 0 ) to excite into resonance only spins contained in a predefined slice. 
   Each of the gradients G x , G y , and G z  then provide for prephasing pulses  38  of a type well known in the art and G x  is controlled to produce a series of echoes  40  represented by echo train  41  whose acquisitions each provide a line of k-space data as will be described. The G x  gradient series produces either a positive frequency encoding pulse  42  or a negative frequency encoding pulse  44  used to generate the echo train  41 . As will be understood in the art, the positive frequency encoding pulse  42  provides for a left to right traversal of k-space (in the k x  direction) whereas the negative frequency encoding pulse  44  provides a right to left traversal of k-space eliminating the need for unproductive rewinding in k-space. 
   In between each echo  40  of the echo train  41 , the G y  gradient provides a phase encoding pulse  46  and  48  which moves the path of k-space traversal in the k y  direction downward for positive phase encoding pulses  46  and upward for negative phase encoding pulses  48 . 
   Referring now to  FIGS. 2 and 3 , for the first four lines of k-space labeled: a, b, c and d, positive phase encoding pulses  46   a - 46   c  are of equal area causing three equal steps through k-space downward in the k y  direction after each k-space line of data is acquired. Subsequent positive phase encoding pulse  46   d , however, has twice the area of positive phase encoding pulses  46   a - 46   c  causing the next line of k-space acquired to be line f skipping line e. 
   The next phase encoding pulse is a negative phase encoding pulse  48   a  equal in area magnitude to the area magnitude of positive phase encoding pulses  46   a - 46   c  (but negative in sign) to cause a backtracking or retracing in the k y  direction of the trajectory of acquisition back to line e after line f has been acquired. 
   This negative phase encoding pulse  48   a  is followed by positive phase encoding pulse  46   e  equal in area to positive phase encoding pulse  46   d  and similarly reinstituting a forward motion along k y  in k-space to k-space line g. 
   A next negative phase encoding pulse  48   b  retraces in the k y  direction to rescan line f previously scanned. Thus line f is scanned twice in opposite directions separated by a predetermined interval in time. 
   This negative phase encoding pulse  48   b  is followed by positive phase encoding pulse  46   f  equal in area to positive phase encoding pulse  46   d  and similarly reinstituting a forward motion along k y  in k-space to k-space line h. A similar process as described above causes line g to be scanned twice in opposite directions separated by the same predetermined interval in time as the two scans of line f. Lines i, j and k are then scanned in succession without multiple scanning. 
   Thus lines f and g are both scanned twice and the remaining lines scanned once. The phase of the image data evolves in time from the moment of the RF excitation pulse  34  and is proportional to magnetic field strength as derived from the Bloch equations. The Bloch equation is well known in the art and described in Magnetic Resonance Imaging: Physical Principles and Sequence Design, E. Mark Haacke, Robert W. Brown, Michael R. Thompson and Ramesh Venkatesan, Editors. John Wiley &amp; Sons, New York (1999) hereby incorporated by reference. Thus, the phase difference of images derived from the k-space data of lines f and g reveal the magnetic field strength over the area of the image and variations in that field strength reveals inhomogeneities. 
   Referring now to  FIG. 4 , in an alternative embodiment, lines e and f, for example, may be acquired twice successively without the intervening acquisition of other k-space lines. In this case, given positive phase encoding pulses  46  are simply suppressed to cause a retracing in k-space. In this case, the time separation between redundant acquisitions is less, with the result that sensitivity is somewhat reduced but the danger of phase “roll-over” in which phase information becomes ambiguous, is reduced. 
   Note that in both cases, a full plane of k-space image data is acquired for the purpose of constructing an image and the doubly acquired data for the purpose of generating the magnetic field map is limited to a few lines. These lines are preferentially center k-space lines so as to provide for the important low spatial frequencies in the magnetic field map. High spatial frequency information for the magnetic field map preferably not acquired so as to limit the increase of the acquisition time, but this is optional and the amount of high frequency data acquired may be flexibly adjusted according to acquisition speed requirements. 
   Referring now to  FIG. 5 , generally many single lines of k-space data  52  covering an entire plane will be obtained including edge k-space image data  54   a  and  54   b  providing high frequency image data and center k-space image data  54   c  providing low frequency image data. Duplicated or redundant lines of k-space magnetic field data  56  providing low frequency magnetic field map data will preferentially but optionally cover only a few lines near the center of k-space. 
   For each application of the pulse sequence per  FIG. 2 , the center k-space image data  54   c  and the duplicated lines of k-space magnetic field data  56  are summed by summing block  58  (implemented by program  28 ) to make full use of the additional image data in the duplicated lines of k-space magnetic field data  56  to increase the signal-to-noise ratio of this k-space data. The output of summing block  58  is then combined with edge k-space image data  54   a  and  54   b  and transformed by a two-dimensional Fourier transform of transform block  59  to produce an image  60  for each echo  40 . The pulse sequence of  FIG. 2 , having a single RF excitation pulse  34  producing a echo train  41  with multiple echo  40  thus produce the data for a single image  60  and its corresponding magnetic field map  64 . Repeated applications of the pulse sequence of  FIG. 2  over time result in a time course  67  of images  60  over a period such as may provide the basis for a functional magnetic resonance imaging (fMRI) study as will be understood to those of ordinary skill in the art and an example of which will be described below. The images  60  may be displayed on the console  32  and further manipulated as is understood in the art. 
   The duplicated lines of k-space magnetic field data  56  and center k-space image data  54   c  are then reconstructed using a two dimensional Fourier transform per blocks  61  and  63  or other technique and the resulting phase information analyzed by phase analyzer  64  per the Bloch equation to produce a magnetic field map  65  for each echo  40 . The same repeated applications of the pulse sequence of  FIG. 2  over time used to generate the time course  67  of images  60  also result in a time course of magnetic field maps  64 , each of which provides a snapshot of the changing in homogeneities in the magnetic field of the MRI machine  10 . 
   Referring now to  FIG. 6 , the program  28  operating in the controller  26  of the MRI machine  10  provides in an RF excitation pulse  34  per process block  70  after which multiple lines of k-space data as described above may be acquired as indicated by process block  72 . 
   As described above, a magnetic field map is extracted from the doubly sampled k-space data near the center of k-space as indicated by process block  74  while image data are extracted from the combined doubly sampled and singly sampled k-space data as indicated by process block  76 . 
   In one embodiment, each magnetic field map  64  of process block  74  may be used as indicated by arrow  78  to correct the corresponding image  60  of process block  76  (that is the image derived from the same data acquired at the same time) as indicated by process block  80 . Each of these corrected images may be added to a time series study as indicated by process block  82 . Alternatively, at process block  80 , images  60  that have more than a predefined distortion from magnetic field inhomogenities detected at process block  74  may be discarded. 
   Referring to  FIG. 5 , generally the correction on an image  60  is effected by phase shifting the underlying k-space data of the image  60  according to the inhomogeneity measured by the corresponding magnetic field map  64 . This correction, in one embodiment, may be applied only within a zone defined by a predetermined high intensity portion of the image  60  which approximates the portion of the image occupied by the subject. Note that the phase correction must be modified according to the time separation between the acquisitions of the k-space lines which is no longer uniform because of the double sampling. Thus, generally phase corrections in the k y  direction are shifted by half as much in the outer lines of k-space than the central lines when double sampling occurred. 
   In a second embodiment, each magnetic field map  64  of process block  74  may be used as indicated by process blocks  84  and  86  to correct the magnet shimming for the next image  60  (that is the image derived from the data acquired after the next RF excitation pulse  34 ). This may be done using the most recent magnetic field map  64  only or from an extrapolation of the magnetic field map  64  and previous magnetic field maps  64 . 
   As shown in  FIG. 5 , this extrapolation may, for example, fit a curve to a value of corresponding pixels  66  in several previous magnetic field maps  64 , to deduce an extrapolated value for the corresponding pixel  68  of a next magnetic field map  64 ′. This extrapolation, repeated for each pixel  66  in the previous magnetic field maps  64 , predicts the next magnetic field map  64 ′ and may be used to adjust the shimming coils of the magnet  12  through shimming subsystem  22  for the next image  60 ′. 
   Dynamic shimming is made possible by the present invention which allows both magnetic field map information and image data to be obtained using a single RF pulse, at little penalty, on a near real-time basis. 
   Through the geometric correction of process block  80  and the shimming of process block  86 , either of which may be performed alone or together, distortion caused by magnetic inhomogeneity and the long readout times of echo planar imaging and other similar techniques are substantially reduced. 
   Referring now to  FIG. 7 , the increase in acquisition time caused by the double sampling of the present invention may be reduced by combining the present invention with a technique of acquiring approximately half of the k-space data  88  and filling the missing k-space data in with the Hermitian conjugate formation  90  described in detail in “Single-Shot Half k-Space High-Resolution Gradient-Recalled EPI For fMRI At 3 Tesla” by Andrzej Jesmanowicz, Peter A. Bandettini and James S. Hyde, Magnetic Resonance in Medicine 40:754-762 (1998), hereby incorporated by reference. In this case, the center lines of k-space may be acquired two more times after the full k-space trajectory has been completed and used to correct mismatches that can occur between forward and backward k x  lines of k-space. 
   Alternatively, the partial k-space acquisitions can be used to provide higher spatial resolution of the image as described in the above-cited paper. More generally, the above partial k-space acquisition allows flexibility in trading acquisition time against image resolution in the context of the present invention. 
   The invention may find use not only in fMRI studies, but also in imaging of ostensibly static anatomical structures where involuntary patient motion nevertheless may be a concern. Here the echo planar gradient recalled pulse sequence can be tailored to enforce a relatively short acquisition window to reduce blurring of the image during involuntary motion at the expense of signal-to-noise ratio. The echo planar gradient recalled pulse sequence is intrinsically fast and may be further increased in speed, if desired, through the use of the partial k-space acquisition technique described above. 
   Motion of the patient during multiple acquisitions of time series  67  thus resolves itself as a series of shifted images  60 , each with low motion blur, but with the patient captured in different relative positions. Combining these images  60 , to improve the signal-to-noise ratio, can be done in theory by correcting for the motion (adjusting the relative position of the patient in each image by means of rigid body translations and rotation) and then averaging the corresponding pixels of the images, on a pixel by pixel basis. Motion correction of this type, relying on an analysis of correlation between the images with varying amounts of rotation and translation, is well known in the art. 
   As a practical matter, however, precise motion correction required to generate a high signal-to-noise composite image is hampered by the susceptibility distortion introduced into the image data when a patient moves within a magnetic polarizing field shimmed for a different patient position. This susceptibility distortion, which is not uniform over the image, cannot be corrected by rigid body displacement of the image (typical motion correction) and thus hampers precise alignment of the images needed for their combination. Misalignment effectively reintroduces blur into the composite image. 
   Accordingly, referring to  FIG. 6 , the present invention further contemplates that at process block  82 , the susceptibility distortion corrected images from process block  80  (optionally as corrected by dynamic shimming of process block  86 ), are motion corrected then averaged to produce a sharp, high signal-to-noise ratio composite image. 
   EXAMPLE I 
   As described briefly above, each application of RF excitation pulse  34  results in an image  60 . In this context, the RF excitation pulse  34  is sometimes referred to as “single shot”. Each of the echoes  40  in the echo train  41  produces a line of k-space. A representative time to acquire an image  60  is 100 milliseconds. A time TR, typically 2 seconds, is defined as the time between acquisitions of an image  60  from a particular slice, however, within TR, multiple images  60  of different parallel slices, for example  20 , can be acquired allowing image data to be acquired from an extended volume of tissue. This is termed “multislice acquisition”. 
   In an fMRI study, to which the present invention may be advantageously applied, data are acquired by imaging the entire human brain using 20 slices 5 millimeters thick where each slice is imaged 128 times over the course of a study. Thus a time equal to 128×TR=256 seconds is typically required to collect the fMRI data. 
   An “image time course”  67  of images  60  (shown in  FIG. 5 ) is defined as consisting of 128 images  60  from a single slice. Typically there may be  20  such image time courses  67  to cover the entire brain. It follows that a typical whole brain fMRI data set consists of 20×128=2560 images  60 . Each of these 2560 images requires application of the pulse sequence shown in  FIG. 2 . 
   According to the present invention, 2560 magnetic field maps  64  are also produced in a typical FMRI experiment corresponding to the 2560 images. There are 20 magnetic field map time courses  65  (shown in  FIG. 5 ) in this example corresponding to the 20 image time courses  67 . The 20 magnetic field map  64  acquired within each TR provide a mapping of the complete brain. However the 20 magnetic field map  64  are not perfectly registered in time with each other since the underlying data of each magnetic field map  64  is acquired, typically, displaced in time by 100 ms. Such a time course of whole brain magnetic field maps do, however, characterize magnetic field variation in time that occur slower than the TR value. 
   Each volume element (“voxel”) of tissue is sampled in this example 128 times. A voxel of tissue gives rise to a pixel  66  describing signal intensity in an image. Thus each voxel gives rise to a “pixel time course” of image intensity. A “voxel magnetic field time course” can be produced from the magnetic field map time course data that corresponds to each pixel time course. 
   The present invention is particularly useful for fMRI studies in which jaw and tongue motion may be present, for example, where the study requires speaking by the subject. Here, the susceptibility changes caused by anatomical motion may be corrected to better reveal the blood oxygen level-dependent (BOLD) effect or similar fMRI signals. It is specifically intended that the present invention not be limited to the embodiments and illustrations contained herein, but include modified forms of those embodiments including portions of the embodiments and combinations of elements of different embodiments as come within the scope of the following claims.