Rapid MRI using multiple receivers producing multiply phase-encoded data derived from a single NMR response

Abstract: Method and apparatus for more rapidly capturing MRI data by receiving and recording NMR RF responses in plural substantially independent RF signal receiving and processing channels during the occurrence of an NMR RF response. The resulting plural data sets respectively provided by the plural RF channels are then used to produce multiply phase-encoded MRI data from the single NMR RF response. Practical examples are disclosed for reducing required MRI data capturing time by factors of at least about one-half.

Description: SPECIFICATION This invention is generally related to magnetic resonance imaging (MRI) using nuclear magnetic resonance (NMR) phenomena. It is particularly directed to method and apparatus for more efficiently capturing and providing MRI data suitable for use in multi-dimensional Fourier transformation MRI imaging processes. MRI is by now a widely accepted and commercially viable technique for obtaining digitized video images representative of internal body structures. There are many commercially available approaches and there have been numerous publications describing these and other approaches to MRI. Many of these use multi-dimensional Fourier transformation techniques which are, by now, well-known to those skilled in this art. For example, in one commercially available MRI system, a slice selective G z magnetic gradient pulse is utilized in conjunction with RF nutation pulses (including a 180° nutation pulse) to produce true spin echo NMR RF pulses from only a relatively narrow "planar" or "slice" volume perpendicular to the Z axis. During readout and recordation of the NMR spin echo RF response, a G x magnetic gradient pulse is employed to provide spatially dependent frequency/phase-encoding in the X axis dimension. Accordingly, by a first one-dimensional Fourier Transformation process, one can obtain Fourier coefficients representing the NMR spin echo response at different X locations emanating from a correspondingly located "column" volumes parallel to the Y axis. By rapidly repeating this same process using different slice selective G z gradient pulses during a single T1 NMR interval, it is known that one can significantly enhance the efficiency of obtaining data for a number of planar volumes (sometimes called "multi-slice" MRI). However, one cannot produce the requisite two-dimensional visual image from only a single dimension of Fourier transformation (per slice) as just described. To obtain the second dimension of Fourier coefficients for resolution in the Y axis dimension, a different G y phase-encoding pulse (e.g., different in magnitude and/or time duration) is utilized during the NMR excitation process such that the NMR spin echo responses during different data gathering cycles will produce Fourier coefficients phase-encoded with respect to spatial location in the Y axis dimension. Accordingly, if one wants to obtain, for example, a resolution of 256 pixels along the Y axis dimension, then one must go through 256 data gathering cycles (per slice) with correspondingly different Y axis phase-encoding in each cycle to assemble the requisite data required for the second dimension of Fourier transformation. Because the normally encountered T1 NMR parameter is on the order of a second or so in many human tissues, and because one typically does not (in this exemplary system) repeat a data taking cycle within the same volume until the previous NMR nuclei have substantially returned to their quiescent conditions, it will be appreciated that the need to repeat a data taking cycle literally hundreds of times translates into an overall MRI data capturing process that requires several minutes to complete. And some of the more interesting future MRI applications may even require thousands of data taking cycles and several tens of minutes using current technology. At the same time, as is well known in the art, MRI systems, facilities and operating personnel represent a significant expense which can only be economic if the time required for capturing NMR data per patient is minimized. In addition to the rather basic economic improvement that can be obtained by reducing MRI data capture time, it may also make practical some rather sophisticated new MRI possibilities. For example, three dimensional MRI may be useful for permitting one to obtain oblique reconstruction images on any desired (oblique) plane. For this application, one should use isotropic resolution which implies the creation of an MRI data set having 256 voxels (volume picture elements) in each of three mutually orthogonally dimensions. Even if partial flip imaging techniques are used to minimize intervals between NMR excitation sequences, current MRI techniques might still require imaging times on the order of a half hour -- which is probably beyond a reasonable imaging time for most patients. If this time requirement could be reduced by at least a factor of one-half, it might become a much more attractive possibility. Similarly, echo planar imaging techniques (which permit all data required for a single image to be obtained after a single excitation) would become more practical if higher resolution data could be obtained during the ensuing train of NMR RF responses (which is necessarily limited by the T2 decay parameter). The enhanced diagnostic possibilities that may someday be provided by spectroscopic MRI imaging might also become more practical if some technique is developed for shortening the time required to collect requisite spectroscopic imaging data. MRI angiography is another technique which presently requires unusually long MRI imaging times and which may become considerably more practical if a reduced imaging time technique could be employed. Furthermore, it is perhaps self-evident that motion artifact can be reduced if one can somehow shorten the time interval over which requisite MRI data is collected. In short, standard techniques of two-dimensional Fourier transform magnetic resonance imaging are already highly efficient in two out of three dimensions. Through the use of selective excitation and data acquisition with a readout gradient, the Fourier transform of the acquired data effectively localizes signal in two dimensions. However, there remains a time consuming aspect of MRI insofar as it is still necessary to obtain multiple phase-encoded data acquisitions to localize the signal in the third (Y axis) dimension in many standard MRI processes. The need for reducing MRI image time requirements has been recognized by many others. There are other techniques for more rapidly obtaining requisite MRI data. For example, other techniques for rapid MRI imaging have attempted to reduce the time spent on acquiring phase-encoded acquisitions. For example, data conjugation techniques (exploiting the symmetries of the Fourier transform) are already employed in some commercial processes to reduce the number of required acquisitions by a factor of two. Partial flip angle nutation (instead of a full 90° initial RF nutation angle) may also be employed to eliminate signal loss normally encountered with rapid acquisitions (e.g., as in short repetition time techniques). In principle, the most efficient means of signal acquisition for MRI may involve the use of only partial initial RF nutation angles or "partial flip" combined with a very short TR interval between repetitions of the NMR excitation-response processes used in the data gathering phase of MRI. For example, this latter partial flip technique has aided in the development of three dimensional Fourier transform techniques for imaging a large number of sections within a very short TR. Some publications generally relevant to such MRI techniques as have just been discussed may be noted as follows: Kumar A, Welti D and Ernst RR. NMR Fourier Zeugmatography. Journal of Mangetic Resonance 1975; 18:69-83. Sutherland RJ and Hutchison JMS. Three-Dimensional NMR Imaging Using Selective Excitation. Journal of Physics E 1978; 11:79-83. Lauterbur PC. Image Formation By Induced Local Interactions: Examples Employing Nuclear Magnetic Resonance Nature 1973; 16:242-243. den Boef JH, van Vijen CMT and Holzscherer CD. Multiple-Slice NMR Imaging by Three-Dimensional Fourier Zeugmatography. Physics in Medicine and Biology 1984; 29:857-867. Feinberg DA, Hale JD, Watts JC. Kaufman L and Mark A. Halving MR Imaging Time By Conjugation: Demonstration at 3.5KG. Radiology 1986; 161:527-531. Ernst RR. Sensitivity Enhancement in Mangetic Rescnance In: Waugh JS, ed. Advances in Mangetic Resonance Vol. 2, New York: Academic Press, 1966:1-135. Carlson J, Crooks LE, Ortendahl DA, Kramer DM and Kaufman L. Technical Note: Comparing S/N and Section Thickness in 2-D and 3-DFT MRI. Radiology 1988;166:266-270 Some aspects of my invention have already been published ("An Algorithm for NMR Imaging Reconstruction Based on Multiple RF Receiver Coils," Journal of Magnetic Resonance Vol. 74. Sept. 1987. pages 376-380). This prior art is reflected in the preambles of claims 1 and 8. And others, subsequent to my invention, have now proposed the general concept of using multiple detectors for MRI data acquisition so as to save data acquisition time -- but without any apparent practical implementations in mind (see Hutchinson et al, "Fast MRI Data Acquisition Using Multiple Detectors," Magnetic Resonance in Medicine, Vol. 6, 1988, pp 87-91). This invention is especially useful in MRI data capturing sequences where a dimension transverse to the static magnetic field is phase-encoded over a plurality of data gathering sequences to obtain requisite MRI data for a single image. It may be used with both two and three-dimensional Fourier Transform Magnetic Resonance Imaging processes. Preferably, in the exemplary embodiments, the static magnetic field is horizontal (as in a super conducting solenoidal magnet) so that the exemplary RF receiving coils permit the most convenient patient access. According to the invention there is provided a method for magnetic resonance imaging the internal structure of an object to produce an image having resolution MxN pixels in x and y dimensions within a planar volume of substantially fixed z dimensions, MRI data for said image being captured using NMR RF responses of nuclei subjected to a static magnetic field with a sequence of mutually orthogonal x,y,z magnetic gradient pulses and NMR RF excitation pulses superimposed thereon, comprising the steps of: (i) selectively exciting a planar volume of NMR nuclei of substantially fixed z dimension using a z-axis oriented magnetic gradient; (ii) phase-encoding nuclei in a y-axis dimension using a subsequent pulse of y-axis oriented magnetic gradient; (iii) in the presence of an x-axis oriented magnetic gradient, (a) recording a resultant NMR RF response from said nuclei via a first RF receiving coil and RF signal processing channel to obtain first MRI data, and (b) simultaneously also recording said resultant NMR RF response from said nuclei via at least a second RF receiving coil and RF signal processing channel to obtain second MRI data; (c) each receiving coil comprising a plurality of elongated conductors disposed parallel to the static magnetic field and providing an independent NMR RF response even though all receiver coils are coupled to a common imaging volume; (iv) processing said first and second MRI data to provide further MRI data representative of at least two different degrees of y-axis phase-encoding, at least one of which degrees of y-axis phase-encoding is different from the degree of phase-encoding actually employed in step (ii); and (v) repeating steps (i)-(iv) substantially less than N times, using different degrees of y-axis magnetic gradient, to provide a complete MRI data set capable of producing an MRI image having N pixel resolution along the y-axis dimension; said method being characterized by step (v) using y-axis magnetic gradients which, on different repeat cycles, provide at least two sets of said further MRI data representing the same degree of y-axis phase encoding; and step (v) also including the step of averaging together said further MRI data representing the same degree of y-axis phase encoding. Furthermore, there is provided an apparatus for magnetic resonance imaging the internal structure of an object to produce an image having resolution MxN pixels in x and y dimensions within a planar volume of substantially fixed z dimensions, MRI data for said image being captured using NMR RF responses of nuclei subjected to a static magnetic field with a sequence of mutually orthogonal x,y,z magnetic gradient pulses and NMR RF excitation pulses superimposed thereon, comprising: (i) means for selectively exciting a planar volume of NMR nuclei of substantially fixed z dimension using a z-axis oriented magnetic gradient; (ii) means for phase encoding nuclei in a y-axis dimension using a subsequent pulse of y-axis oriented magnetic gradient; (iii) means for, in the presence of an x-axis oriented magnetic gradient, (a) recording a resultant NMR RF response from said nuclei via a first RF receiving coil and RF signal processing channel to obtain first MRI data, and (b) simultaneously also recording said resultant NMR RF response from said nuclei via at least a second RF receiving coil and RF signal processing channel to obtain second MRI data; (c) each receiving coil comprising a plurality of elongated conductors disposed parallel to the static magnetic field and providing an independent NMR RF response even though all receiver coils are coupled to a common imaging volume; (iv) means for processing said first and second MRI data to provide further MRI data representative of at least two different degrees of y-axis phase encoding, at least one of which degrees of y-axis phase-encoding is different from the degree of phase-encoding actually employed in step (ii); and (v) means for repeating said exciting, phase encoding and recording substantially less than N times, using different degrees of y-axis magnetic gradient, to provide a complete MRI data set capable of producing an MRI image having N pixel resolution along the y-axis dimension; said apparatus being characterized by said means for repeating using y-axis magnetic gradients which, on different repeat cycles, provides at least two sets of said further MRI data representing the same degree of y-axis phase encoding; and said means for repeating also includes means for averaging together said further MRI data representing the same degree of y-axis phase encoding. These as well as other objects, advantages and features of the present invention will become more completely understood and appreciated by carefully reading the following detailed description of the presently preferred exemplary embodiments in conjunction with the attached drawings, of which: FIGURE 1 is an abbreviated schematic block diagram of an MRI system employing this invention; FIGURES 2A through 2D are schematic depictions of various birdcage and saddle coil constructions that might be utilized for achieving a practical RF receive coil for use with this invention; FIGURE 3 is a graph showing spin density of a simple one-dimensional sample; FIGURE 4 is a graph depicting standard Fourier reconstruction of the samples shown in FIGURE 3; FIGURE 5 depicts a two-coil reconstruction of a point sample using three projections per echo; FIGURE 6 depicts a two-coil reconstruction of a point sample using two projections per echo; FIGURE 7 depicts a two-coil reconstruction of a point sample using an averaged value of two projections per echo; FIGURE 8 depicts a standard Fourier reconstruction of the same sample with Gaussian white noise added to the sample signal; and FIGURE 9 depicts a two-coil reconstruction of the same sample with sample generated noise. In FIGURE 1 an MRI system 100 is schematically depicted as including the usual static magnet, gradient coils, shim coils, transmit RF coils, 102 under control of processor 104 (which typically communicates with an operator via a conventional keyboard/control display module 106). Although, it is conceivable that a single processor might both control the system and also carry out the actual MRI imaging processes, it is perhaps more conventional to employ a system of multiple processors for carrying out specialized functions within the MRI system 100 as will be appreciated. Accordingly, as depicted in FIGURE 1, an MRI image processor 108 receives digitized data representing RF NMR responses from an object under examination (e.g., a human body 110) and, typically via multiple Fourier transformation processes well-known in the art, calculates a digitized visual image (e.g., a two-dimensional array of picture elements or pixels, each of which may have different gradations of gray values or color values, or the like) which may then be conventionally displayed at 112. In accordance with this invention, a plurality of receive coils 1...N are independently coupled to a common imaging volume (e.g., a desired portion of body 110). The RF signals emanating from these coils are respectively processed in independent RF channels 1 through N. As depicted in FIGURE 1, each such RF channel may comprise a considerable amount of conventional analog RF signal processing circuits as well as an eventual analog to digital conversion before being input to the MRI processor 108 (which may typically include means for digitally storing the acquired data during a data acquisition sequence until the image processor 108 uses such acquired data to produce an image at 112). Although the RF channel circuitry may typically include a rather complex (and expensive) amount of circuitry, it is not believed necessary to describe it in any detail since conventional RF signal processing circuitry per se may be employed with this invention. However, the extra expense of using extra receive coils and associated RF signal processing channels does have to be balanced against the improvement in data capturing time when considering the overall economics of this approach. Accordingly, to practice this invention, modifications need be made in essentially only three areas of a conventional MRI processing system: 1. additional receive coils need to be employed and their construction is preferably such that they are substantially independent (i.e., effectively without substantial mutual inductive coupling or capacitive coupling); 2. an additional RF signal processing channel needs to be added for each of the additonal receive coils; and 3. the MRI reconstruction algorithm programmed into and implemented by the image processor 108 needs to be slightly modified so as to use the additional incoming data to calculate multiply phase-encoded MRI data in an appropriate way. Since the RF signal processing channels are, per se , simply replications of existing conventional RF channels, it is not believed that any further detailed description of such additional channels need be given in this application. Rather, the following disclosure will concentrate on practical exemplary embodiments for the receive coil and for the new calculations to be made in the image processor 108. As will be appreciated by those in the art, practical coil constructions for RF receiving in MRI typically include tuning and impedance matching capacitances in conjunction with transmission lines, etc. For purposes of simplifying the discussion in this case, such conventional aspects of RF receiving coils and their associated RF transmission lines, etc., are not depicted in the FIGURES or otherwise described. Another way to understand the underlying theory, is to imagine a simplified imaging experiment which attempts to reconstruct the transverse magnetization density in a sample tube that is parallel to the y axis. The x and z location of the column are known. What is needed, is a reconstruction of the sample along the y direction. This simplified experiment can be viewed as merely part of a standard two-dimensional Fourier Transformation MRI imaging sequence. A slice/selective RF excitation provides a section of sample with a known value of z and the read-out gradient G x provides frequency encoding in the x-direction. Thus, after Fourier Transformation of the acquired data, the Fourier content of the data in a standard two-dimensional Fourier transform acquisition provides the sum of the transverse magnetization in the hypothetical column or sample now being considered. Standard two-dimensional FT imaging uses repeated acquisitions with phase-encoded echoes to generate the Fourier transform of the transverse magnetization along the y direction. If the field of view along the y direction is taken to be L, the Fourier transform is usually written as m n = ∫ m y e 2π iny / NL dy and m = m x + im y is the complex transverse magnetization. For the purposes of illustration, consider the idealized RF receiver coil for this experiment. The receiver consists of a collection of long wires parallel to the static field and laid on the surface of a cylinder. The loops are closed by return paths at infinity. In this idealized receiver coil one may measure the echo signal by monitoring the voltage induced in each wire. Denote the voltage in the wire at position  by V(). As shown above, the Fourier transform of the magnetization density is expressible in terms of the Fourier transform of the induced voltages. Let V(n) be the Fourier coefficient of the voltage, then the expressions for the Fourier transform of the magnetization take the form: m (0) = ∫ m ( y ) dy = iRV (-1) m +1 = ∫ m y e i π y / R dy = iRe π x / R [ V -1 + i π V -2 - π 2 2 V -3 +···] m (+2) = ∫m ( y ) e 2 iπy/R dy = iRe 2 πx/R [ V (-1)+2 iπV (-2)-2π 2 V (-3) +···] m -1 = ∫ m y e -i π y / R dy = iRe - π x / R [ V -1 - i π V -2 - π 2 2 V -3 +···] m (-2) = ∫ m ( y ) e -2 iπy/R dy = iRe -2 πx/R [ V (-1)-2 iπV (-2)-2π 2 V (-3)+ ···] where R is the radius of the coil; the field of view, L, is 2R. Only the five lowest Fourier coefficients are shown. In principle, given sufficiently many voltage Fourier coefficients, one may construct all terms in the Fourier transform of the magnetization density. Thus, one can reconstruct the entire column of the sample within one spin echo response. There are two important practical considerations that rule out the possibility of a complete reconstruction from one echo, however. One problem is coil coupling. An actual coil of separate loops surrounding a cylinder would not behave exactly as described. One reason is that coils in receivers are tuned circuits and voltages produced by echoes induce currents. Since the idealized receiver coil has a collection of loops with large mutual inductance between loops, the signal in each loop is not that just produced by the echo. The crucial observation - and this is key to feasibility - is that even though one cannot build the idealized coil, one can build separate coils that directly measure the voltage Fourier coefficients where there is substantially no intrinsic mutual inductance between these coils. One embodiment of a multiple coil set is a pair of birdcage resonators. Previously, these coils have been developed as a means of producing a homogeneous RF field. Homogeneity is due to the current distribution on the legs of a driven birdcage that varies as cos . One can use the birdcage as a quadrature transmitter and generate a complex current distribution exp(-i). In this situation, a driving voltage produces a voltage distribution across the legs of the birdcage that varies as exp(-i). By using the standard reciprocity principle, we can see how to produce a coil to measure V(-1). Since the voltage distributions are the same in transmission and reception, a standard birdcage resonator produces a net output voltage proportional to the sum of voltages induced on each leg weighted by exp(-i). A coil that measures V(-2) requires a birdcage coil with a current distribution that varies as cos 2. By extending the previous results one can see that this coil measures a sum of voltages induced on each let with a weighting factor of exp(-2i). One method of producing a cos 2 distribution is to construct a birdcage coil and tune it to its second resonant frequency. Normally, a birdcage coil has many resonant frequencies, each of which is approximately a multiple of the lowest fundamental frequency. What one here requires is one coil whose fundamental frequency coincides with the NMR frequency (e.g., 15 MHz) and a separate coil whose second harmonic is at the same frequency. The two coils may be placed concentrically around the sample with substantially no intrinsic mutual inductance. Since there is substantially no intrinsic coupling between the coils, each coil behaves independent of the other and the net signal on each output is the proper voltage Fourier coefficient. One still has to contend with capacitive coupling between the coils, as well as residual inductive coupling. However, careful adjustment of the relative alignment of the two coils and other now standard techniques for decoupling quadrature-connected receive coils should provide adequate isolation. Birdcage resonators are not the only way to construct these coils. Some other configurations are shown in FIGURES 2A-2D. Note that quadrature as well as non-quadrature coils may be constructed. Non-quadrature coils will be much easier to construct, but the lower signal to noise ratio of these coils may degrade the image substantially. FIGURE 2A depicts a "low pass" birdcage. FIGURE 2B depicts a homogeneous quadrature saddle coil. Arrows denote the direction of current flow. Since the coil has a homogeneous reception pattern, the signal is essentially the same (albeit with higher signal to noise ratio) as a non-quadrature saddle coil (FIGURE 2C). Another design for a coil to measure V(-2) is shown in FIGURE 5D. The second practical consideration facing complete reconstruction from one echo is the signal-to-noise ratio. One expects that the signal to noise ratio of human MRI is insufficient to allow for single echo reconstruction. Therefore, one may concentrate on a reduced implementation. For example, only two voltage Fourier coefficients will be measured in the present exemplary embodiments: V(-1) and V(-2). Given these, one can construct three Fourier coefficients of the magnetization by truncating the sums for the expression of m(n): m (-1) = iRe -πx/R [ V (-1)- iπV (-2)] m (0) = iRV (-1) m (+1) = iRe πx/R [ V (-1)+ iπV (-2)] An actual imaging sequence will be a hybrid technique of reconstruction using standard phase-encoding in the y direction and using the extra spatial information generated by the two voltage signals. That is, a typical data acquisition sequence may first acquire an echo with no phase encoding. From this, one can construct m(-1), m(O) and m(1). Next, a sequence will acquire an echo with a phase-encoding factor of exp(2iπy/R). From this, one can construct the terms m(1), m(2) and m(3). This is repeated for the rest of the positive and all negative phase encodings: Where: M = Fourier coefficients along y-axis dimension M(0) = m a (0) M(1) = [m a (1) + m b (1)]/2 M(2) = m b (2) M(3) = [m b (3) + m c (3)]/2 M(4) = m c (4) . . . M(-1) = [m a (-1) + m q (-1)]/2 M(-2) = m q (-2) M(-3) = [m q (-3) + m r (-3)]/2 M(-4) = m r (-4) . . The net result is that one only requires one-half of the usual number of echoes in order to fully reconstruct the image. Notice that one acquires the even projections (m(O), m(2), m(4), and so on) entirely conventionally from the signal V(-1) from coil #1. The odd projections then are calculated from both V(-1) and V(-2). The odd projections are calculated twice and averaged in the presently preferred embodiment. This provides a more artifact free reconstruction. Examples of one dimensional reconstructions are discussed below. Notice that the coil to measure V(-1) is a uniform receiver coil and resembles standard designs in imaging and spectroscopy. For present purposes, this will be referred to as the "primary" channel. The coil to measure V(-2) is highly non-uniform. This will be referred to as the "secondary" channel. The signal-to-noise ratio of multi (e.g., two) channel reconstruction deserves special treatment. Since, for the case N=2, one has a system of two receivers, the noise in a reconstructed image depends on whether the noise in the two channels is correlated or uncorrelated. If the coils themselves generate the noise, then the noise is uncorrelated in the two channels. Noise generated by the sample produces a noise voltage which is received by both coils and is therefore correlated. (Notice that the noise is correlated but not necessarily in phase in both channels. The relative phase of the noise voltage depends on the location of the noise generating sample.) Consider a specific example. A point sample is placed at the center of the coil. In this simplified arrangement, the magnetization produced by the sample will not generate any signal in the secondary channel. So also, any noise generated by the sample will not be received by this channel. The primary channel, will receive both signal and noise generated by the sample. A quick calculation of the noise in an image with standard phase-encoded imaging and multiple receiver coil imaging shows: Noise Multiple Coil Noise Standard =√2cos 2 π y / L The noise at the center of the image increased by 2. This is just a consequence of the multiple coil acquisition acquiring one-half the data. Therefore, multiple coil acquisition has the same signal-to-noise ratio per unit time as standard imaging in this example. Away from the center of the image, noise actually decreases as compared to standard imaging. Since the noise is correlated in the frequency domain, the Fourier transform of the data (i.e., the image) shows structured noise. This analysis changes if the noise is completely generated by the coils. Using the same example as before, the secondary channel will have a noise voltage even though there is no sample generated signal. Following in the standard noise analysis, the ratio of the noise in two imaging procedures is: Noise Multiple Coil Noise Standard = 2cos 4 π y / L + π 2 2 sin 2 2π y / L The noise has again increased by a factor of 2 at the center. And, unlike the previous example, noise never decreases compared to standard imaging. MRI of biological tissue in mid to high field strengths is in the regime of sample generated noise. There may be a loss in signal to noise ratio. This procedure is applicable to sequences well above the signal to noise ratio threshold where speed is the issue. The expressions for the Fourier transform of the magnetization density, m(n), involve infinite sums of voltage Fourier coefficients, V(n). In practice, as proposed above, one only measures signal from a finite number of coils (e.g., two coils). The effects of truncating the infinite sum at only two terms, for example, is considered below. Also, given signal from only two receiver coils, some Fourier projections of the magnetization density can be calculated more reliably than others. That is, the previous section contains the expressions for m(-1), m(O) and m(1) and V(-2). For the most artifact free reconstruction, calculating all three m(n)'s from one echo (resulting in a 67% time reduction), calculating two m(n)'s from one echo, or calculating three m(n)'s from one echo but redundantly calculating every other m(n) and averaging the results will provide different quality MRI images. Considering these effects for quadrature birdcage coil receivers, in a typical imaging scenario, the acquired time domain signal is Fourier Transfcrmed to yield terms in the Fourier Transform along the y direction for a given x. The remaining problem is a one-dimensional reconstruction of density along the y direction. The profile of such a simple point sample is shown in FIGURE 3. Here, signal only arises from a small sample located off center. A standard 128 projection reconstruction with a single receiver coil is shown in FIGURE 4. This displays the magnitude of the reconstructed profile; the point spread is the usual behavior of a finitely sampled Fourier transform. FIGURE 5 shows the magnitude of a reconstruction using two receivers and calculating three Fourier projections per echo. The artifact in the reconstruction is quite large. A less ambitious approach of calculating two Fourier projections per echo (FIGURE 6) with a 50% time savings decreases the artifact but it is still large. The best results so far are obtained with a redundant calculation of three projections per echo; every other echo is calculated twice and averaged (as noted above in Table I). A magnitude reconstruction with this approach is shown in FIGURE 7. The level of artifact is significantly reduced. The height of the false peak is roughly 2% of the height of the sample peak. The 50% time savings is preserved. FIGURE 8 shows a standard reconstruction of a point sample with sample generated noise. An image acquired in half the time using two receivers shows an increase in noise as shown in FIGURE 9. As expected, the noise level has increased by 2. Towards the edge of the field of view noise decreases. While only a few exemplary embodiments have been described in detail above, those skilled in the art will recognize that many variations and modifications may be made in these exemplary embodiments while yet retaining many of the novel features and advantages of this invention. All such variations and modifications are intended to be encompassed by the scope of the appended claims.

Claims: A method for magnetic resonance imaging the internal structure of an object (110) to produce an image having resolution MxN pixels in x and y dimensions within a planar volume of substantially fixed z dimensions, MRI data for said image being captured using NMR RF responses of nuclei subjected to a static magnetic field with a sequence of mutually orthogonal x,y,z magnetic gradient pulses and NMR RF excitation pulses superimposed thereon, comprising the steps of: (i) selectively exciting a planar volume of NMR nuclei of substantially fixed z dimension using a z-axis oriented magnetic gradient; (ii) phase-encoding nuclei in a y-axis dimension using a subsequent pulse of y-axis oriented magnetic gradient; (iii) in the presence of an x-axis oriented magnetic gradient, (a) recording a resultant NMR RF response from a said nuclei via a first RF receiving coil and RF signal processing channel to obtain first MRI data, and (b) simultaneously also recording said resultant NMR RF response from said nuclei via at least a second RF receiving coil and RF signal processing channel to obtain second MRI data; (c) each receiving coil comprising a plurality of elongated conductors disposed parallel to the static magnetic field and providing an independent NMR RF response even though all receiver coils are coupled to a common imaging volume; (iv) processing said first and second MRI data to provide further MRI data representative of at least two different degrees of y-axis phase encoding, at least one of which degrees of y-axis phase-encoding is different from the degree of phase-encoding actually employed in step (ii); and (v) repeating steps (i)-(iv) substantially less than N times, using different degrees of y-axis magnetic gradient, to provide a complete MRI data set capable of producing an MRI image having N pixel resolution along the y-axis dimension; said method being characterized by step (v) using y-axis magnetic gradients which, on different repeat cycles, provide at least two sets of said further MRI data representing the same degree of y-axis phase encoding; and step (v) also including the step of averaging together said further MRI data representing the same degree of y-axis phase encoding. A method as in claim 1, characterized in that steps (i)-(iv) are repeated approximately N/2 times. A method as in claim 1, characterized in that step (iv) comprises: (a) obtaining the first Fourier Transform coefficient V(1) of the first MRI data; (b) obtaining the first Fourier Transform coefficient V(2) of the second MRI data; (c) calculating three Fourier coefficients m(-1), m(0) and m(+1) of NMR magnetization for respectively corresponding different degrees of y-axis phase encoding using the formulae: m(-1) = iRexp(-πx/R)[V(1)-iπV(2)] m(0) = iRv(1) m(+1) = iRexp(πx/R)[V(1)+iπV(2)] where R = the effective radius of the RF receiving coils, and x = the x coordinate for which a column of y-axis phase encoded nuclei is being analyzed. A method as in claim 3, characterized in that the m(-1) result for one repeat of steps (i)-(iv) is averaged with the m(+1) result for another repeat of steps (i)-(iv). A method as in claim 1, characterized in that said first and second RF coils used in step (iii) each comprise birdcage coils, said first RF coil being resonant at a first fundamental frequency of about F and said second RF coil being resonant at a second fundamental frequency of about F/2. A method as in claim 1, characterized in that said first and second RF coils used in step (iii) each comprise a pair of QD-connected saddle coils (1, 2), said first RF coil being resonant at a first fundamental frequency of about F and said second RF coil being resonant at a second fundamental frequency of about F/2. A method as in claim 5 or 6, characterized in that said first RF coil provides a current/voltage distribution which varies as sin  about said image volume,  being an angle measured about the coil center relative to a predetermined elongated conductor ; and said second RF coil provides a current/voltage distribution which varies as sin 2 about said image volume. An apparatus for magnetic resonance imaging the internal structure of an object (110) to produce an image having resolution MxN pixels in x and y dimensions within a planar volume of substantially fixed z dimensions, MRI data for said image being captured using NMR RF responses of nuclei subjected to a static magnetic field with a sequence of mutually orthogonal x,y,z magnetic gradient pulses and NMR RF excitation pulses superimposed thereon, said apparatus comprising: (i) means (102, 104) for selectively exciting a planar volume of NMR nuclei of substantially fixed z dimension using a z-axis oriented magnetic gradient; (ii) means for phase-encoding nuclei in a y-axis dimension using a subsequent pulse of y-axis oriented magnetic gradient; (iii) means (102, 104, 108) for, in the presence of an x-axis oriented magnetic gradient, (a) recording a resultant NMR RF response from said nuclei via a first RF receiving coil and RF signal processing channel to obtain first MRI data, and (b) simultaneously also recording said resultant NMR RF response from said nuclei via at least a second RF receiving coil and RF signal processing channel to obtain second MRI data; (c) each receiving coil comprising a plurality of elongated conductors disposed parallel to the static field and providing an independent NMR RF response even though all receiver coils are coupled to a common imaging volume; (iv) means (108) for processing said first and second MRI data to provide further MRI data representative of at least two different degrees of y-axis phase encoding, at least one of which degrees of y-axis phase-encoding is different from the degree of phase-encoding actually employed in step (ii); and (v) means (104) for repeating said exciting, phase encoding and recording substantially less than N times, using different degrees of y-axis magnetic gradient, to provide a complete MRI data set capable of producing an MRI image having N pixel resolution along the y-axis dimension; said apparatus being characterized by said means (104) for repeating using y-axis magnetic gradients which, on different repeat cycles, provides at least two sets of said further MRI data representing the same degree of y-axis phase encoding; and said means (104) for repeating also including means for averaging together said further MRI data representing the same degree of y-axis phase encoding. An apparatus as in claim 8, characterized in that said means (104) for repeating causes approximately N/2 repetitions to occur. An apparatus as in claim 8, characterized in that said means for processing comprises (a) means for obtaining the first Fourier Transform coefficient V(1) of the first MRI data; (b) means for obtaining the first Fourier Transform coefficient V(2) of the second MRI data; (c) means for calculating three Fourier coefficients m(-1), m(0) and m(+1) of NMR magnetization for respectively corresponding different degrees of y-axis phase-encoding using the formulae: m(-1) = iRexp(-πx/R)[V(1)-iπV(2)] m(0) = iRv(1) m(+1) = iRexp(πx/R)[V(1)+iπV(2)] where R = the effective radius of the RF receiving coils, and x = the x coordinate for which a column of y-axis phase encoded nuclei is being analyzed. An apparatus as in claim 10, characterized in that said means for averaging averages the m(-1) result for one repeat with the m(+1) result for another repeat. An apparatus as in claim 8, characterized in that said fist and second RF coils each comprise birdcage coils, said first RF coil being resonant at a first fundamental frequency of about F and said second RF coil being resonant at a second fundamental frequency of about F/2. An apparatus as in claim 8, characterized in that said fist and second RF coils each comprise a pair of QD-connected saddle coils (1, 2), said first RF coil being resonant at a first fundamental frequency of about F and said second RF coil being resonant at a second fundamental frequency of about F/2. An apparatus as in claim 12 or 13, characterized in that said first RF coil provides a current/voltage distribution which varies as sin  about said image volume,  being an angle measured about the coil center relative to a predetermined elongated conductor; and said second RF coil provides a current/voltage distribution which varies as sin 2 about said image volume.

Inventors: CARLSON JOSEPH W, CARLSON, JOSEPH W.
Applicants: UNIV CALIFORNIA, THE REGENTS OF THE UNIVERSITY OF CALIFORNIA
Application year: 1989
Main classification: G01R33
IPC codes: G01R33, A61B5