Patent Publication Number: US-11385312-B2

Title: Ultrafast MRI system and method

Description:
REFERENCE TO RELATED APPLICATIONS 
     This patent application is a continuation—in part of parent application Ser. No. 16/166,709 filed Oct. 22, 2018, which in turn is a continuation-in-part of Ser. No. 14/671,528, now U.S. Pat. No. 10,107,883, and incorporates by reference both said earlier-filed patent applications. 
    
    
     FIELD 
     This patent specification is in the field of imaging, such as medical imaging, and more specifically relates to rapid three-dimensional (3D) magnetic resonance imaging (MRI). 
     BACKGROUND 
     Several references are identified by numerals in parenthesis in this patent specification and are hereby incorporated by reference. The full citations are listed at the end of the specification. References [18] and [19] are papers by the inventor and persons working under his direction related to the subject matter of the paper. 
     Real-time MR imaging could exert a profound influence on neuroscience in the future by enabling the direct visualization of neuronal interactions. For now, however, the known practical embodiments of MRI require at least some degree of gradient encoding, and this in turn sets a lower limit of about 100 ms for volume acquisition. 
     In the original formulation of MRI by Lauterbur [1], spatial encoding is achieved by applying successive magnetic field gradients to the imaging volume. Each new gradient is associated with a different radiofrequency (RF) excitation, and RF-induced echoes form a line in k-space, discretized into N elements, where N is the dimension of the image matrix. After N progressively increasing gradients, and N echoes, there are N lines in k-space, and the N×N k-space matrix is subjected to a 2-D Fourier transform, rendering the N×N image matrix. In echoplanar imaging (EPI) Mansfield [2] showed that spatial encoding could be achieved by trains of gradient reversals after a single RF excitation. An advantage of using gradient reversals is speed; a disadvantage is sensitivity to susceptibility changes and magnetic field inhomogeneities. 
     More recently studied parallel MRI (pMRI) uses the spatial sensitivities of multiple receiver coils (detectors) arranged around the object for spatial encoding of voxels in a single slice, thus reducing the need for gradient reversals and RF pulses [3, 4]. pMRI images are discussed by Kelton, Magin and Wright [5], and by Ra and Rim [6]. pMRI modifies the Lauterbur-Mansfield approach to include multiple receiver coils, so that if the number of coils is n, then the number of gradient reversals is a factor n times smaller. According to such studies of pMRI, many of the conventional 180° RF pulses can be replaced by short trains of gradient reversals, with an acceptable change in image quality. Thus, pMRI can embrace some of the advantages of EPI, without some of the disadvantages. Because gradient switching is generally several times faster than RF excitations, image quality can be adequately maintained while speed is increased somewhat. 
     The pMRI initially proposed by Hutchinson [3,4] is for single shot imaging that ignores sources of noise in the surface coils. Subsequent analyses of detector arrays by Roemer [7] and Ocali [8] account to some extent for the effects of noise, with algorithms for conventional gradient encoding. A concept of “ultimate signal-to-noise,”[8] can be considered, by which was meant that, at least for gradient encoding, large numbers of small receiver coils—when properly configured—can be more efficient than a single receiver. Sodickson [9], with SMASH, and Pruessmann [10], with SENSE, proposed merging gradient-encoding and spatial sensitivity encoding. These merging techniques can be designated “hybrid” since they use both methods of encoding. 
     Merging of the two fundamental ways of encoding comes at a considerable price, because for the hybrid techniques ultimate signal-to-noise is reached not with large numbers of coils, but with small numbers (typically 4). Signal-to-noise (SNR) in hybrid pMRI is reduced by 3 sources: (i) electrical noise, (ii) reduced numbers of echoes, and (iii) a critical geometric factor, g, as discussed by Ohliger [11] and Wiesiger [12]. Sources of electrical noise include preamplifiers, coupling between adjacent coils, and body thermal noise. In addition, there is noise from the eddy currents in the body surface due not only to the gradient reversals but also to re-radiation of signal by the coil, which is both a receiver and a dipole radiator. These sources of noise can be reduced, however, and for conventional gradient encoding multiple small coils have been shown to be more efficient than one large coil [7,8] so there is no insurmountable noise problem inherent to small coils. When multiple coils are used to encode, however, there can be an additional problem due to the inefficiency of spatial sensitivity encoding, compared with gradient encoding, when the coils are large. This in turn is due to the slow spatial variations of the spatial sensitivity of each coil. In addition, since there are fewer phase-encoding steps SNR is further reduced. The signal-to-noise ratio is now given by [11,12]:
 
SNR=SNR Full /(√ R ) g   (1)
 
where SNR Full  is the SNR for gradient-only encoding, R is the acceleration factor (in this case the number of detectors) and g is the geometric factor. The loss factor JR represents the reduced SNR due to the reduced number of phase encoding steps and could not be mitigated absent phase encoding. The g-factor is intrinsic to the geometry of the coil array and is a measure of the capacity of the array to compensate for reductions in gradient encoding. For very small numbers of detectors g is close to 1, but as the number of coils increases, and the number of phase-encoding steps correspondingly decreases, g suddenly becomes large and the images are degraded.
 
     It has generally been accepted that for practical purposes this single limitation means that for hybrid techniques acceleration factors need to be at most about 4, so that the rate of acquisition of SNR is about double that of non-parallel techniques. The reason g&gt;1 with increasing accelerations is that the transformations used to obtain the image in conventional pMRI are non-unitary, and the reason for this is that the solutions to Maxwell&#39;s equations are “smooth,” by which is meant that for large detectors the spatial dependence of the radio field is not as granular as the spatial dependence of the gradients used for conventional encoding. This in turn means that large groups of adjacent pixels may have similar spatial sensitivity profiles. 
     As R increases, and the number of gradient-encoding steps correspondingly decreases, more and more of the burden of encoding is shouldered by the receiver array. For R= N  n=4, only 25% of the encoding is gradient based. With more detectors there is an unsupportable reliance on spatial sensitivity encoding and this leads to rapid image degradation when R&gt;4. The spatial sensitivity maps can be modeled for illustrative purposes by taking, as an approximation to the solutions of Maxwell&#39;s equations, only the 1/r 3  dependence of the near field dipole. This view of an inherent limitation of all constructions of parallel MRI has been challenged by Keil and Wald [13], with a theoretical evaluation suggesting that the spatial sensitivity of small coils could contribute more in spatial encoding in hybrid pMRI while still using some gradient encoding. The original concept for single shot, single slice imaging free of gradient reversals [3,4] was tested several years ago by McDougall and Wright [14] using a 64-channel coil with an acceleration factor of 64. However, this was not for volume encoding but only for a single slice. Proposals have been published to employ static magnetic field gradients produced by thin magnetic films to encode flow [15], to use two or three RF phase gradients in an arrangement free of magnetic gradients [16], and to use a flexible array coils populated with multiple coils in multi-slice-multi-echo sequences that inherently rely on magnetic field gradients [17]. 
     The prior commercial embodiments of pMRI known to the inventor herein require a plurality of magnetic gradients and/or gradient reversals for spatial encoding. This can generate a large amount of noise, first RF and other electrical noise, and second audible noise. This means that image signal-to-noise ratio is lowered, and that the machine can be extremely loud (up to 120 dB). 
     This patent specification recognizes and addresses these and other aspects of prior MRI work by providing a new, radically different approach. 
     SUMMARY OF THE DISCLOSURE 
     This patent specification describes a new approach using an MRI system that spatially encodes a volume with a multitude of very small receiver coils in a 3D array around the volume. In the presence of a steady magnetic gradient g z  and a steady main field B 0  in an imaging volume, an initial excitation RF pulse creates an echo that contains the information needed for a 3D image of the entire imaging volume. (See coordinate system defined in  FIGS. 1 a  and 1 b   .) 
     RF pulses subsequent to the initial excitation pulse can be applied to the imaging volume to refocus and this create additional echoes. In addition, or instead, gradient reversals can be applied to the imaging volume to create echoes subsequent to the echo from the initial excitation RF pulse. The subsequent RF pulses can reverse inherent inhomogeneities in the steady magnetic gradient g z  as well as correct for inherent inhomogeneities in the main magnetic field B 0 . The gradient reversals can reverse inhomogeneities in the gradient field g z  though not in the main field B 0 . 
     Notably, spatial encoding in XY planes is provided by the 3D array or small coils around the imaging volume. Any subsequent RF pulses and/or gradient reversals are used to refocus and this create echoes like the echo from the initial excitation pulse rather than echoes for different XY spatial encodings. This is unlike conventional MRI in which gradient pulse sequences must provide spatial encoding in XY planes. 
     One example of the new approach comprises a gantry with a magnet configured to generate a main magnetic field B 0  in an imaging volume, a gradient field generator configured to generate a steady gradient field g z  in the imaging volume, a radio-frequency (RF) pulse generator configured to apply an initial excitation RF pulse to the imaging volume, and a multitude of small MR signal receiving coils arranged in a three-dimensional (3D) array surrounding the imaging volume and extending along the B 0  field. Each of the receiver coils is configured to simultaneously receive RF energy from the entire imaging volume during MR signal acquisition and output respective MR signals. An MR signal acquisition facility acquires the MR signals and a computer-implemented processor applies image reconstruction algorithms to the MR signals and thereby generates a 3D image of an object in the imaging volume, displayed as such or as two-dimensional (2D) images derived therefrom. The MR signal from each of the coils comprises an initial echo from the excitation pulse that contains information for reconstructing a 3D image of the object. If additional RF pulses and/or gradient reversals are applied, a time sequence of subsequent overall peaks and valleys follows the initial MR signal and can extend over a period of the order of seconds and even minutes without requiring additional spatial encoding 
     The 3D array of receiving coils comprises M coils along the direction of the B 0  field and N coils transverse to B 0 . Spatial encoding is based on spatial sensitivity parameters of the receiving coils. The gradient field generator is configured to generate the steady gradient field g z  as a single gradient field maintained throughout the MR signal acquisition (or possibly reversed one or more times to create additional echoes), and the acquisition facility is configured to acquire the MR signals in the substantial absence of gradient field reversals or additional RF pulses for XY plane encoding, although inhomogeneities in the B 0  field can be compensated with additional RF pulses if used for refocusing and thus creating echoed subsequent to the echo from the initial excitation pulse. The image reconstruction can involve applying a one-dimensional Fourier Transform or Fourier Series process to the MR signal from the respective coils to generate transformed MR signals and applying a matrix multiplication process to the transformed MR signals using a matrix related to spatial sensitivities of the receiving coils. 
     The new approach also comprises carrying out an MRI process comprising applying a main magnetic field B 0 , a steady gradient magnetic field g z , and an excitation radio-frequency (RF) pulse to a three-dimensional (3D) imaging volume, acquiring RF energy for the entire imaging volume with each of a multitude of receiving coils arranged in a 3D array around the imaging volume during MR signal acquisition in the substantial absence of encoding in XY planes by gradient reversals or subsequent RF pulses, applying image reconstruction algorithms to the MR signals to thereby generate a 3D image of an object in the imaging volume, and displaying the 3D image as such or as two-dimensional (2D) images derived therefrom. Subsequent RF pulses and/or gradient reversals may be used to refocus and thus create echoes like but subsequent to the echo from the initial excitation pulse. An embodiment of another aspect of the new approach is a computer program product embodied in a non-transitory form in a computer-readable medium and comprising instructions that, when executed by a computer system, cause the system to carry out the steps of acquiring magnetic resonance (MR) signals for an entire imaging volume generated from each of a multitude of receiving coils arranged in a 3D array around an imaging volume during MR signal acquisition in the substantial absence of spatial encoding in XY planes by gradient reversals or RF pulses subsequent to an initial RF pulse, applying image reconstruction algorithms to the MR signals to thereby generate a 3D image of an object in the imaging volume, and displaying the 3D image as such or as two-dimensional (2D) images derived therefrom. Subsequent RF pulses and/or gradient reversals may be used to refocus and thus create echoes like but subsequent to the echo from the initial excitation pulse. Because each of the receiving coils receives MR signals from the entire 3D imaging volume and because of the substantial absence of spatial encoding in XY planes by pulse sequences and in conventional MRI, the output of each of the coils is remarkably superior to the signals that a single receiving coils outputs in a conventional MRI system, or each of the few coils in a known pMRI system outputs. For example, the MR signals from each of the small coils is comparable in quality to the signals received by a single coil in a conventional MRI system. The implications of this include that for an image of quality comparable to that from a conventional MRI system, the new system described in this patent specification can (i) acquire MRI signals over a much shorter period, (ii) use much lower strength main magnetic field, (iii) greatly improve the SNR (signal-to-noise ratio) of the images, and/or (iv) create much higher resolution images for comparable or even lower main magnetic field and/or for shorter MR signals acquisition times. For example, the SNR of MRI signals from the new system can be at least an order of magnitude better than for MR signals obtained under otherwise similar conditions from a currently commercially available 1.5 T MRI system. 
     The benefits of needing much lower main magnetic field include conveniently making the new MRI systems essentially open-MRI, for example by making the magnet opening very large or making the magnet in portions that are spaced apart by several feet or even meters. As an extreme case, the new system can eliminate magnets as a source of the main magnetic field and rely solely on the Earth&#39;s magnetic field as a source of the main magnetic field B 0 . The Earth&#39;s magnetic field can be roughly 10 −4  times the strength of the main magnetic field in a conventional MRI scanner, but the stronger MR signals of the multitude of small coils in the new system can largely compensate for this, assuming appropriate measures are taken to ensure adequate uniformity of the Earth&#39;s magnetic field acting on the imaging volume. 
     In addition, because the new system does not rely on gradient pulse sequences for spatial encoding in XY planes, and because gradient pulse sequences are a highly significant source of electronic noise that causes image degradation in conventional MRI, the MR images from the new system can be much less noisy compared to a conventional MRI system. This is so because much, if not most of the electrical noise in the MR signals in a conventional MRI system is related to gradient switching. Elimination of gradient switching, or near elimination, can reduce the electrical noise to zero or near zero. As a result, the SNR of the new MRI system can be orders of magnitude better than in a conventional MRI system. In addition, complete or substantial elimination of gradient pulses can reduce or eliminate the loud acoustical noise that many patients find troublesome. 
     As an example, if the 3D array of small coils is n×n coils (n rings each of n receiver coils), and each coil is A/n of the area A of the larger receiver coil of a conventional MRI system, the MR signals received by the small coil would have a strength much greater that 1/n 2  of the strength of the MR signals received the larger coil. (“n” is the number of small coils.) This is so because the smaller coil in effect “sees” all the MR signals from all the 3D imaged volume, all the time. In contrast, in conventional MRI systems, each echo encodes for one line in k-space and the signal is correspondingly reduced by a factor of n 2 . 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWING 
         FIG. 1 a    illustrates in block diagram form a portion of an MRI system according to one embodiment described in this patent specification;  FIG. 1 b    illustrates a relevant coordinate system;  FIG. 1 c    illustrates a portion of an array of receiver coils and illustrates in block-diagram form the RF pulse generator, the source of the steady magnetic gradient, an MR signals acquisition unit, the image reconstruction processor, and the image display unit of a complete MRI system;  FIG. 1 d    illustrates in perspective view a portion of the array of the 3D receiver coils; and  FIG. 1 e    illustrates a stack of 2D arrays of receiver coils. 
         FIG. 2  illustrates an array of receiver coils when unfolded and flattened. 
         FIG. 3  illustrates an MR signal generated by a receiver coil from an initial echo and subsequent echoes. 
         FIG. 4  illustrates a prior art arrangement of receiver coils. 
         FIG. 5 a    illustrates an essentially or virtually open-MRI configuration with a very large diameter magnet bore, and  FIG. 5 b    illustrates an essentially or virtually open MRI configuration with a main magnetic field generated by widely spaced apart magnetic components that can be free-standing or integrated in room walls or floor and ceiling. 
     
    
    
     DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS 
     An example of a preferred embodiment can operate by using some of the basic components of a conventional MRI system, such as a magnet generating a main magnetic field B 0  in an imaging volume, a source of RF pulses, a source of a magnetic gradient, a computer configured to process MR signals and to control the overall operation of the system, and a workstation to format and process MR images and allow operator control. The new approach described in this patent specification adds other components, and programs the computer in the system to carry out different operations. The new components include a drastically different 3D array of very small receiver coils surrounding the imaging volume where the output of each of the small coils is an MR signal comprising contributions from the entire imaging volume. The new approach further includes modifications in the control over, and if desired the structure of, the RF pulse and gradient generators, and providing a different facility for computer-processing the output of the multitude of small receiver coils. 
     One example of a preferred embodiment is illustrated in block diagram form in  FIGS. 1 a -1 e    and comprises an MRI gantry  10  that can be the same as or like MRI systems currently offered commercially by companies such as Siemens Healthcare and GE Healthcare. For example, gantry  10  comprises a magnet  102  generating a main magnetic field B 0 , using a superconducting magnet surrounding a cylindrical imaging volume  104 .  FIG. 1 b    shows a Cartesian xyz reference in which the z-direction is along the length of a patient table  110  and the x,y directions are perpendicular to the table length. RF pulse generator  106  applies RF pulses to the imaging volume, and a steady gradient field generator  108  generating a steady magnetic gradient g z . An MR signal acquisition unit  120  acquires echoes (MR signals), and an image reconstruction processor  122  reconstructs MR images and controls the system components and their operation. 
     Unlike conventional MRI systems, gantry  10  includes a 3D array  112  of a multitude of very small receiver coils (detectors)  114 , which array extends in the z-direction and surrounds the imaging volume. Also unlike conventional MRI systems, gantry  10  is configured to generate a steady gradient magnetic field g z  extending in the z-direction that does not require changes for spatial encoding as in a conventional MRI system although optionally a few gradient reversals could be used for a different purpose, namely, to reverse inherent inhomogeneities in the B 0  field and/or to create echoes subsequent to the echo from the initial excitation RF pulse. Still in addition, gantry  10  differs from conventional MRI systems in that its RF pulse generator is configured to generate a single RF excitation pulse, such as a 90° excitation pulse, and need not generate additional RF pulses for XY encoding during MR signal acquisition but can be configured to provide several subsequent RF pulses for a different purpose, namely, to refocus spins and thus create subsequent echoes and/or to compensate for inhomogeneities of the B 0  field. For conciseness, this patent specification does not provide details of the conventional MRI components and processes, so as not to obscure the new components and processes that are described in greater detail below. 
     The 3D array  112  of receiver coils comprises a multitude of very small, loop-shaped receiver coils  114 , each for example of the order of several mm along each side. 3D array  112  can be generally cylindrical in shape, surrounding the imaging volume, although it can deviate somewhat from a circular cross-section, for example to an oval shape or a more rectangular shape with rounded corners. The receiver coils  114  are arranged in generally round 2D arrays  116 , which 2D arrays can be stacked next to each other in the z-direction. The 2D arrays  116  surround the imaging volume. A 3D array  112  can be formed, for example, by a 3D printing technique that embeds the conductive elements of each coil (and related conductive lines to the coils) in a material that is not electrically conductive, such as a polymer, and is tubular or is initially flat but later is rolled into a tubular shape to surround the imaging volume of the MRI system. The number of such 2D arrays  116  is related to the desired length of the 3D imaging volume  104  along the z-direction and the desired spatial resolution in that direction. For example, if the 3D imaging volume  104  is considered as a stack of slices each 1 mm thick in the z-direction, and if the imaging volume is 20 cm long in the z-direction, then there can be a stack of approximately 200 2D arrays  116  of receiver coils  114  forming the 3D array  112 . However, there need not be a one-to-one correspondence between the number of 2D arrays of receiver coils and the number of actual or virtual slices of the imaging volume. 
       FIG. 1 c    illustrates a spatial arrangement of receiver coils  114  in a magnified partial view, where one of the coils is labeled  114   mn  to illustrate an example of “m” 2D arrays  116  of coils  114  and “n” coils  114  in each 2D array  116 . The 2D arrays may all have the same number of coils per array but in another embodiment, some 2D arrays may have different numbers of coils.  FIG. 1 c    also schematically illustrates that the MR signals from coils  114  go to a MR signals acquisition facility or unit  120  and are processed into MR images in image reconstruction processor  122  for display on display  124  (or for other purposes such as computer image analysis or manipulation, storage, etc.).  FIG. 1 c    also illustrates RF generator  106  that produces the required RF pulse or pulses and a gradient generator  108  that maintains a steady magnetic gradient during MR signals acquisition (or if desired provides gradient reversals).  FIG. 1 d    illustrates a stack of several 2D arrays  116  of coils. 
       FIG. 2  illustrates an arrangement of small receiver coils  114  when the 3D array  112  of coils is unfolded and flattened for easier visualization. There can be M 2D arrays  116  of coils  114  in the z-direction, which can but need not correspond to M actual or virtual XY slices of the imaging volume. Each 2D array  116  can comprise N coils  114 , where M×N can be 256×256 in one non-limiting example. M and N are positive integers much greater than unity that may but need not be the same. Each of the coils  114  is identified by the reference numeral  114   mn , to indicate that it is the m-th coil in the z-direction in the 3D array  112 , and the n-th coil within the 2D array  116  of coils  114  in a direction transverse to the z-direction. Each detector  114  is depicted as a single loop. As an example, for a spatial resolution of 1 mm and a field of view of 250 mm, each of the loops  114  can be approximately 3 mm perpendicular to the z-direction. Preferably each coil  114  is less than 1 cm in each dimension; more preferably, each is less than 5 mm in each dimension; still more preferably each coil is 3 mm or less in each dimension. 
     Gantry  10  excites an object in the imaging volume  104  with a single RF pulse, for example a 90° RF pulse, while the object is in the imaging volume and an initial steady gradient field g z  is being applied to the imaging volume. In the example of a 1.5 T MRI magnet, the initial gradient field can have a typical strength of 45 mT/meter. Unlike in conventional MRI, the gradient magnetic field g z  need not change during MR signals acquisition for the purpose of spatial encoding although it can be reversed once or several times for spin refocusing the created additional echoes. In addition, there needs to be only an initial RF excitation pulse for generating MR signals that contain information for an image of the entire 3D object in the imaging field although subsequent RF pulses can be used to create additional echoes and correct from main magnetic field inhomogeneities. 
       FIG. 3  illustrates an MR signal S that a single receiver coil  20  generates in response to the initial excitation RF pulse and subsequent RF pulses and/or gradient reversals. Notably, the response MR signal S does not simply rise and fall, as a conventional MR echo in a conventional MRI system, but rather oscillates, forming repeated overall peaks and valleys. Each peak contains information for the entire imaging volume. This oscillation can go on indefinitely if the T 1  and T 2  parameters of the object in the imaging volume did not decay in amplitude with time. In practice, the oscillations can provide useful MR signal of the order of seconds and minutes consistent with the quantum decay (spontaneous emission) of the excited state. Even without decay and without gradient switching, the amplitude of the MRI signal S of each of the small coils goes to nominal zero every τ=(FOV)/(256γg x ) seconds, where FOV is the field of view, γ is the gyromagnetic ratio, g x  is the magnetic field gradient in the x direction, and 256 is the image matrix dimension in pixels 
     The new process is explained below in greater detail. As noted, it makes use of a massively parallel, 3D array  112  of very small receiver coils (detectors)  114  arranged around the object being imaged. Unlike the previous proposal for a single-slice pMRI (3, 4), each of the small coils  114  in the new approach simultaneously acquires MR signals from the entire 3D imaged volume. This leads not only to markedly reduced imaging times, but also to marked increase in MR signal per unit of time. 
     While in theory current MRI systems could acquire signal from the entire volume simultaneously by successive application of gradients in the x, y and z directions, and could analyze the signal with a 3D Fourier transform, in practice after each successive gradient in the z-direction the entire set of spins would have to be allowed to relax, and then re-excited for further signal acquisition. Otherwise the T 2  dephasing would quickly eliminate the useful MR signal. Furthermore, the required periods of relaxation would greatly extend actual acquisition times. 
     In contrast, in the new system described here, which can be identified by the label ULTRA, there can be a steady, single gradient g z  employed throughout the MR signal acquisition. This single magnetic gradient g z  need not be switched or even varied, although it can be reversed once or several times to thereby create subsequent echoes. As a result, the technique can be essentially noise-free both electrically and acoustically. In practice, to overcome or reduce the effect of inhomogeneities of B 0 , the ULTRA process preferably uses a periodic imposition of inhomogeneity-compensating 180° RF pulses, for example approximately every 10 ms. The resulting electrical noise can be a low-frequency hum at approximately 100 Hz that can be filtered out, if desired, for example with a suitable low-pass filter. 
     The massively parallel 3D array of small coils (detectors)  114 , M×N in number, arranged in a generally cylindrical or round fashion around the object in the imaging volume, also is very different from the receiver in the earlier description of pMRI [3, 4], where only a single slice would be excited at any given time in order to provide sufficient amplitude of the detected MR signal. 3D image acquisition and 3D image reconstruction were not proposed or contemplated. The detector array that was believed to be most consistent with maximal signal in the earlier pMRI proposals [3,4] consisted of N long parallel loops on a cylinder as illustrated in  FIG. 4  (where the plane of the imaged slice is perpendicular to the lengths of the coil loops). 
     In the new ULTRA system, it has been unexpectedly discovered that the MR signal per unit time is very significantly increased, if not maximized, by acquiring MR signals from the entire imaged volume simultaneously at each of the small receiver coils  114 , provided the receiver coils (detectors) are made much smaller than in conventional MRI systems. The use of such much smaller receiver coils  114  according to the new approach paradoxically enables a much higher rate of acquisition of MR signal and/or strength of the acquired signals. 
     For simplicity, and to illustrate the new system and method, first consider a hypothetical case in which the parameters T 1  and T 2  of an object in the imaging volume  104  do not decay with time, so that there is no decay of MR signal with time and the image is defined only by the spin density ρ ijk , where (i,j,k) are the quantized (pixelated) coordinates corresponding to (x,y,z) voxels in an object in the imaging volume. The principal steps in an example of the new process can be:
     1. Apply a single steady gradient g in a fixed direction, say g x , which remains fixed throughout the MR signal acquisition;   2. Excite all spins in the entire volume of the object being imaged, simultaneously, with a 90-degree RF pulse.
 
The MR signal S mn;ijk (t) in a small coil (detector)  114   mn  due to volume element (i,j,k) at time (t) in the object in the imaging volume is then
 
 S   mn;ijk ( t )= R   mn;ijk ρ ijk   e   iωt   (2)
 
where ω=xg x , x is distance in the x-direction, and R is a matrix representing detector spatial sensitivity of coils  114  in 3D array  112 , based on a near-field dipole spatial dependence of 1/r 3 , where r is the distance between a given spin and a given small coil (detector)  114 , and an angular dependence as given by Maxwell&#39;s equations. Such a spatial dependence can be modified at high field strengths and low wavelengths, but this is known in MRI technology and the way to do it is within ordinary skill in that technology.
 
Since ω is proportional to x, it can simply replace x and Eq. (2) can be rewritten:
 
 S   (ω)   mn;jk ( t )= R   (ω)   mn;jk ρ (ω)   jk   e   iωt   (3)
 
where the left hand side is the amplitude of the MR signal in small coil (detector)  114   m,n  due to the volume element at location (j,k) in the yz plane defined by frequency ω (or, in other words, the yz plane defined by fixed x). Then:
 
 S   (ω)   mn ( t )=Σ R   (ω)   mn;jk ρ (ω)   jk   e   iωt   (3)
 
Let the one dimensional Fourier transform of S(t) be    (ω) , where    (ω)  is a matrix representing the ω th  component of the set of Fourier transformed MR signals in all small coils  114  (all detectors).
 
With this definition, Eq. (4) can be rewritten as:
 
ρ (ω)   ={R   (ω) } −1     (ω)   (5)
 
which in words means the following: the two-dimensional matrix representing the spin densities within the plane defined by frequency ω, is the product of the inverse spatial sensitivity matrix of the small coils  114  in 3D array  112  and the matrix representing the Fourier transform of the time dependent signals in all small coils (detectors)  114 .
   

     Eq. (5) above has some similarities to Eq. (15) of reference [4] but explains a process that is not described or suggested in reference [4]. The left-hand side of Eq. (5) above represents a two-dimensional image plane in a process based on a 3-dimensional structure where all image planes are acquired simultaneously. In contrast, the left-hand side of Eq. (15) of reference [4] represents a line of image data in the context of a 2-dimensional structure where all lines in only a single image plane are acquired simultaneously. Thus, Eq. (5) above achieves the results described in this patent specification in the context of the radically different structure of a 3D array  112  of receiver coils  114  each of which is made much smaller than in reference [4], and the entire imaging volume can now be imaged with one excitation. 
     Eq. (5) above embodies a key concept of the new approach and demonstrates that the entire imaging volume, including all planes of equal frequency defined by a single linear and steady gradient g, can be decoded uniquely by first performing a Fourier transform of the MR signal in each small detector  114 , and then multiplying the transformed signals by an inverse of the matrix of spatial sensitivities of the receiver coils  114 . It can now be seen that each small detector loop  114  receives MR signal information from the entire imaging volume. Therefore, all the MR signals, from all the spins, all of the time, are detected by each receiver coil  114  in the entire set of M×N detectors  114 . This vastly increases the rate of signal acquisition and puts ULTRA in a unique category, since in known conventional embodiments of MRI at any given time only a partial volume of spins gives rise to MR signals. It therefore follows that ULTRA is not only faster than known previous proposals and embodiments but also gives rise to much greater rates of MR signal acquisition. Put another way, if the entire set of M×N detectors  114  in ULTRA were joined to form one large conventional detector, this would capture a much greater MR signal than a large conventional detector employing gradient reversals, since all spins in the object volume would be contributing to the signal at all times. Nevertheless, such an arrangement would not make instant 3-dimensional imaging possible, because there is only one coil. ULTRA takes advantage of an inherently very high MS signal by distributing it throughout a very large number of small coils, and then reconstructing it. With reference to  FIG. 2 , where the matrix of receiver coils  114  is mapped on a two-dimensional plane, it now becomes apparent that ULTRA can be considered a holographic technique in some respects. That is, the entire three-dimensional volume of spin density in the imaging volume can be simultaneously and uniquely encoded in a two-dimensional plane (on which the generally cylindrical array of small coils can be unfolded). 
     The MR signal processing described above can be carried out in a processor such as currently used in commercial MRI systems but programmed differently to carry out the processing, using a program that a person of ordinary skill in MRI and computer programming can write without undue experimentation based on the explanation herein. Such a program can become a computer program product when stored in a non-transitory manner in computer-readable media such an optical disc, a thumb drive, or a magnetic disc. 
     Referring again to  FIG. 3 , as noted above the oscillating MR signal would go on forever for infinite T 1  and T 2 , or in practice until the quantum decay (spontaneous emission) of the excited state ends, which can be on the order of many seconds and even minutes. The signal is periodic, and even without decay, and even without gradient switching, its amplitude (vertical axis in  FIG. 3 ) would go to nominal zero every τ=(FOV)/(256γg x ) seconds, where FOV is the field of view, γ is the gyromagnetic ratio, g x  is the magnetic field gradient in the x direction, and 256 is the image matrix dimension. Therefore, in the special, hypothetical case of infinite T 1  and T 2 , the signal is a periodic function such that within each period τ the amplitude would go to nominal zero at least once, as dictated by the steady magnetic gradient. In real life, not only are T 1  and T 2  finite, but for short time scales the MR signal will decay because of inhomogeneities in B 0 . A rough estimate of the latter decay time can be made by assuming an inhomogeneity of p parts per million. In the example of a 1.5 T (Tesla) MRI system, where the frequency is 65 MHz, and for p=1, the signal would begin to decay in a significant way after 1/65 seconds, or approximately 15 ms. To correct for this, assuming longer MR signal acquisition is desired, a periodic inhomogeneity-compensating 180° RF pulse can be applied, say every 10 ms or some other suitable period. For p=2, in this example the RF pulse would be applied approximately every 5 ms, etc. The imposition of such inhomogeneity-compensating RF pulses would lead to a low-frequency humming sound, which would be quite different from the high frequency, high decibel sounds of gradient reversals which characterize current MRI technology and are disturbing to patients. Note that these inhomogeneity-compensating RF pulses are not used to encode, only to re-phase any dephasing caused by inhomogeneity in B 0 . 
     While the foregoing description pertains to spin density images, in clinical MRI embodiments images can be of interest that are sensitive also to the spin-lattice relaxation time, T 1 , and the spin-spin relaxation time, T 2 . Such images also can be obtained using the ULTRA principles explained in this patent specification. Note first that from each broad peak in  FIG. 3  an entire 3-dimensional image can be created. If T 1  and T 2  were infinite, then each broad peak would have the same height (MR signal amplitude). As one example, for 10 peaks at a time, comprising say 10 ms of MR signal, the entire signal can be averaged into one image having high signal-to-noise ratio. The next set of images, from 10-20 ms, can likewise be averaged into one image, and so on, until MR signal has been acquired for, say, 10 high-signal image sets spanning 100 ms. T 2  can then be determined for each pixel within the imaging volume by fitting it to an exponential decay generated from all image sets. 
     Likewise, in this example after 100 ms the spins can be rephased by a partial excitation RF pulse, and the process repeated, giving yet more MR signal for the T 2 -weighted images, while at the same time allowing for an estimation of T 1 , by the relative amplitudes of the broad peaks compared with the amplitudes after the initial 90° RF excitation pulse. The entire process can be repeated every 100 ms, so that after, say, 1000 ms all information pertaining to both T 1  and T 2  can be obtained with high signal-to-noise ratio. Still likewise, similar principles can be used to create inversion-recovery images or diffusion-weighted images. For example, inversion-recovery images can be made modifying the conventional 180° RF pulse, followed after a time TI by a 90° RF pulse, so that the 90° pulse occurs, say, at time TI−50τ, with acquisition beginning at that time, and continuing until TI+50τ. Diffusion-weighted images might be obtained in analogy with conventional imaging, by applying diffusion gradients every z milliseconds. 
       FIG. 5 a    illustrates an MRI system according to the principles described herein that is essentially or virtually an open-MRI system. Because of the high strength and quality of the MR signals in the new approach described above, the main magnetic field can be much lower strength than in conventions MRI system, and can be provided as one example by a magnet  102  that has a diameter so much greater than of the bore of a conventional MRI system that a patient is much less likely perceive it as a confining, tunnel-like enclosure. For example, the magnet may be spaced from the imaging volume by 3 feet or more in each direction, so that for a cylindrical magnet the distance from the inner surface of the bore to the imaging volume is 3 feet or more, preferably is 4 feet or more, and even more preferably is 5 feet or more. While the main magnetic field applied to the imaging volume can be much less than in a current conventional MRI system, the resulting image can be comparable or better because of the MR signal characteristic in the new system described above. For further patient comfort and/or to allow observation of the patient during MR signal acquisition, the 3D array of coils  114  can be mounted on a transparent or at least translucent structure surrounding the patient. In other respects, the arrangement of  FIG. 5 a    can be structured and can operate as the ULTRA system described above, with elements  106 ,  108 ,  120 ,  122 , and  124 , plus any additional facilities for further image processing and/or image storage. As one example, the main magnetic field can be at least an order of magnitude less than 1.5 T at the imaging volume, where the imaging volume is sufficiently large to encompass a cross-section of the torso of a human patient. 
     Similarly,  FIG. 5 b    illustrates an MRI system that a patient may perceive as essentially or virtually open-MRI system because the main magnetic field is provided by a magnet in the form of two structures  102   a  and  102   b  that are spaced apart by several feet or meters. For example, the distance from the imaging volume to the nearest surface of the magnet is 3 feet or more, preferably is 4 feet or more, and even more preferably is 5 feet or more. One or both of structures  102   a  and  102   b  can be built into a wall or walls of the room, or one of the structures or both structures can be free-standing. In all other respects, the arrangement of  FIG. 5 b    is as in  FIG. 5 a   , i.e., it includes elements  106 , 0108 , 120 ,  122 , and  124  plus any additional facilities for further image processing and/or storage. 
     In case the Earth&#39;s magnetic field is relied on as the source of the main magnetic field, the new system can eliminate the magnet  102  altogether and can be as illustrated in  FIG. 1 d   —the entire MRI system can be the 3D array  112  of coils  114 , RF generator  106  and gradient generator  108  acting on the imaging volume, signal acquisition facility  120 , image processor  122 , and image display  124  (plus any desired additions such as workstations for further image processing and any image storage facilities). 
     While several embodiments are described, the new subject matter described in this patent specification is not limited to any one embodiment or combination of embodiments described herein, but instead encompasses numerous alternatives, modifications, and equivalents. In addition, while numerous specific details are set forth in the description to provide a thorough understanding, some embodiments can be practiced without some or all of these details. Moreover, for clarity, certain technical material that is known in the related art has not been described in detail in order to avoid unnecessarily obscuring the new subject matter described herein. It should be clear that individual features of one or several of the specific embodiments described herein can be used in combination with features or other described embodiments. Further, like reference numbers and designations in the various drawings indicate like elements. Accordingly, the present embodiments are to be considered as illustrative and not restrictive, and the body of work described herein is not to be limited to the details given herein, which may be modified within the scope and equivalents of the appended claims. 
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