A magnetic resonance imaging method and system

Abstract: In a magnetic resonance imaging system, a read gradient scaler (102) scales the amplitude and width of read gradients (78, 82, 84) . A sampling rate control (104) controls the sampling rate of the resonance signals received from corresponding magnetic resonance echoes (80, 54, 86) . For example, when the amplitude of the gradient pulse is doubled and its width halved, the sampling rate of the resultant magnetic resonance signal is doubled, e.g., from a bandwidth of 3210MHz to a bandwidth of 64 MHz. In this manner, some echoes are read-out over a longer period of time with a lower bandwidth to produce lower signal-to-noise data lines; whereas, other echoes are much shorter and are read-out more quickly, but with a lower signal-to-noise ratio. The phase-encoding gradients (74, 88, 90) are controlled (70) such that the resultant data lines with the lowest bandwidth have the lowest phase-encoding angles closest to the zero phase-encoding and the data lines with the highest bandwidth have the highest phase-encoding angles closest to the negative and positive maximum phase-encoding gradients. The data lines are sorted (110) by phase encode angle and bandwidth and reconstructed (120) into an image representation.

Description: The present invention relates to a magnetic resonance imaging method and system. It finds particular application in conjunction with echo-planar (EPI) and gradient and spin echo (GSE) imaging techniques and will be described with particular reference thereto. However, it is to be appreciated that the present technique is also applicable to other rapid imaging sequences including spin echo and field echo, as well as a multiple-echo sequences. Heretofore, subjects have been positioned in a temporally constant magnetic field such that selected dipoles preferentially align with the magnetic field. Radio frequency signals have been applied to cause the preferentially aligned dipoles to resonate and emit resonance signals of a characteristic radio frequency. The radio frequency magnetic resonance signals from the resonating dipoles are read out for reconstruction into an image representation. To strengthen the magnetic resonance signals, the resonance signal is commonly refocused into an echo. A spin echo is generated by following a radio frequency excitation pulse with a 180° refocusing pulse which causes the resonating spin system to refocus as a spin echo. The time between the refocusing pulse and the spin echo is the same as the time between the excitation pulse and the refocusing pulse. Other disturbances to the spin system can also be used to induce an echo. For example, reversing the polarity of the magnetic field, particularly the read gradient of the magnetic field, induces a field or gradient echo. Various techniques have been developed for causing a plurality of echoes sequentially following a single excitation. The echoes may include a series of spin echoes, a series of field echoes, or a mixture of field and spin echoes. See, for example, US-A-4,833,407. Traditionally, spatially-encoded magnetic resonance data for image reconstruction uses the same data sampling rate or bandwidth for every one of a multiplicity of views. The analog-to-digital converter is designed and controlled such that it only converts analog resonance signals (either before or after demodulation) with a fixed frequency for each of 256x256 pixel images, 512x512 pixel images, etc. In reconstruction backprojection, each view of magnetic resonance data is collected during a spin or field echo in the presence of a read gradient that frequency encodes spatial position within the object in the direction of the gradient. A series of like acquisitions are repeated, rotating the gradient, but maintaining its amplitude constant with a constant data sampling rate. In two-dimensional Fourier transform imaging, the read gradient is held stationary rather than rotating, but is preceded by a phase-encode gradient pulse in a direction orthogonal to the read gradient. Although the amplitude of the phase encode gradient was stepped to adjust the phase encoding from view to view, the magnitude of the read-out gradient was held constant with a constant data sampling rate. The scheme of data lines with different phase-encodings are typically denoted as k-space data. That is, in k-space, the data line with zero phase-encoding generally extends across the center of k-space. Data lines with the phase-encoding gradient stepped in progressive positive steps are generally depicted as being above the centerline of k-space, and data lines with progressive negative phase-encode steps are depicted as below the centerline of k-space. Typically, k-space has 256, 512, etc. data lines. The data lines are typically sampled the same 256, 512, etc. times to make a square k-space matrix that is Fourier transformed into a square image. To speed up the data acquisition, segmented k-space schemes have been developed in which each excitation is followed by a plurality of data acquisitions. For example, a series of field echoes are generated by rapidly reversing the read-out gradient, e.g., an oscillating read-out gradient of constant amplitude, to produce a series of data lines in k-space following each excitation or shot. If the oscillating read gradient is fast enough, an entire image of data can be acquired with a single excitation, i.e., a single shot technique. Alternately, as illustrated in the above-referenced US-A-4,833,407, each excitation can be followed by a smaller plural number of data acquisition intervals corresponding to a combination of field and spin echoes. The spin and field echoes are collected with a common bandwidth or sampling rate, but are sorted into different segments of k-space to improve image quality. For example, the first echo following each excitation can be placed in the central portion of k-space which contributes more strongly to the resultant image than the data taken at the maximum and minimum phase encode gradients. US-A-4,952,876 proposes the collection of spin echoes in a multi-spin-echo pulse sequence with different bandwidths to permit chemical shift artefacts to be reduced in a first echo by using a stronger gradient with corresponding increased bandwidth. The present invention provides a method of magnetic resonance imaging in which magnetic resonance is excited in selected dipoles, and magnetic resonance echoes are induced, the induced magnetic resonance echoes are read-out along a first axis in the presence of a read gradient, the magnetic resonance echoes are digitized and demodulated to form digital data lines, the data lines are stored and reconstructed by a reconstruction algorithm to generate an image representation, which includes the steps of reading-out a first of the magnetic resonance echoes with a first sampling rate in the presence of a first read gradient of a first amplitude to generate a first digital data line with a first bandwidth and reading-out a second of magnetic resonance echoes with a second sampling rate different from the first sampling rate in the presence of a second read gradient of a second amplitude different from the first read-out gradient amplitude to generate a second digital data line of a second bandwidth different from the first bandwidth, the digital data lines of the first and second bandwidths being reconstructed into one resultant image representation, characterised in that the data line with lower bandwidth is phase-encoded with a smaller phase-encoding angle and the data line with the higher bandwidth is phase-encoded with a greater positive or negative phase-encoding angle. The present invention also provides a magnetic resonance imaging system which includes a magnet for generating a temporally constant magnetic field through an examination region, gradient magnetic field coils and a gradient magnetic field controller for generating at least phase and read magnetic field gradient pulses in orthogonal directions across the examination region, a radio frequency pulse controller and transmitter for inducing dipoles in the examination region to resonance such that radio frequency resonance signals are generated, a receiver for receiving and demodulating the radio frequency magnetic resonance signals to produce a series of data lines, the receiver being associated with an analog-to-digital converter such that the data lines are digital, a data memory for storing the digital data lines, and a reconstruction processor for reconstructing the data lines from the data memory into an image representation, including a read gradient scaler for scaling amplitude and width of selected read gradient pulses such that a bandwidth of the radio frequency magnetic resonance signals received during the scaled read gradient is altered; and a sampling rate control for controlling a sampling rate of each received radio frequency magnetic resonance signal in accordance with the change in bandwidth, wherein the gradient magnetic field and radio frequency pulse controllers are controlled such that data lines with the lowest bandwidth are phase-encoded with the smallest phase-encoding angles and data lines with the greatest positive and negative phase-encoding angles have the highest bandwidth. By placing the lower bandwidth data at the centre of k-space, the image is disproportionately weighted in favour of the lower noise data. Advantageously, the first and a third read gradients have a common amplitude greater than the amplitude of the second read gradient, and the first and a third sampling rates are the same and higher than the second sampling rate such that the first and a third bandwidths are the same and higher than the second bandwidth. This is particularly valuable when the first and third echoes are gradient echoes and the second echo is a spin echo. Particularly, for narrow bandwidth data acquisitions, the phase error resulting from the local magnetic field inhomogeneity is more significant between different echoes of a repeated echo sequence. Although spin echoes are relatively insensitive to field inhomogeneities, field echoes are affected by field inhomogeneities which results in a T 2 * decay of the signal. The greater the time interval between the spin and field echoes, the greater the phase mismatch between the field and spin echoes due to the T 2 * decay. In general, the narrower the read bandwidth of the echoes, the more dramatic is the effect of field inhomogeneities. Moreover, the negative effect increases in severity as additional field and spin echo sets are collected following each of a series of 180E inversion pulses. Although post-processing methods exist to correct for some of these phase differences, the methods are not completely reliable and tend to introduce some loss in resolution in the final image. The wider first and third read bandwidths help avoid the effect of the field inhomogeneity on gradient echoes. One advantage of the present invention is that it opens a whole new avenue of design flexibility for magnetic resonance imaging sequence optimization. Another advantage of the present invention is that T 2 * related phase errors in multiple-echo imaging sequences can be reduced. Another advantage of the present invention resides in the accelerated data acquisition rate. The invention will now be further described by way of example with reference to the accompanying drawings in which: FIGURE 1 is a diagrammatic illustration of a magnetic resonance imaging system in accordance with the present invention; FIGURE 2 is a diagrammatic illustration of an exemplary magnetic resonance imaging sequence including spin and field echoes in accordance with the present invention; FIGURE 3 illustrates a segmentation of k-space in accordance with the present invention; and, FIGURE 4 illustrates an alternate magnetic resonance sequence embodiment for echo-planar imaging. With reference to FIGURE 1, a main magnetic field control 10 controls superconducting or resistive magnets 12 such that a substantially uniform, temporally constant magnetic field is created along a z-axis through an examination region 14. A magnetic resonance echo means applies a series of radio frequency (RF) and magnetic field gradient pulses to invert or excite magnetic spins, induce magnetic resonance, refocus magnetic resonance, manipulate magnetic resonance, spatially and otherwise encode the magnetic resonance, to saturate spins, and the like to generate magnetic resonance imaging and spectroscopy sequences. More specifically, gradient pulse amplifiers 20 apply current pulses to selected ones or pairs of whole body gradient coils 22 to create magnetic field gradients along x, y, and z-axes of the examination region 14 . A digital radio frequency transmitter 24 transmits radio frequency pulses or pulse packets to a whole body RF coil 26 to transmit RF pulses into the examination region. A typical radio frequency pulse is composed of a packet of immediately contiguous pulse segments of short duration which taken together achieve a selected magnetic resonance manipulation. The RF pulses are used to saturate, excite resonance, invert magnetization, refocus resonance, or manipulate resonance in selected portions of the examination region. For whole body applications, the resonance signals are commonly picked up by the whole body RF coil 26 . For generating images of limited regions of the subject, local coils are commonly placed contiguous to the selected region. For example, an insertable head coil 30 is inserted surrounding a selected brain region at the isocenter of the bore. The insertable head coil optionally includes local gradient coils 32 which receive current pulses from the gradient amplifiers 20 to create magnetic field gradients along x, y, and z-axes in the examination region within the head coil. A local radio frequency coil 34 is used to excite magnetic resonance and receive magnetic resonance signals emanating from the patient's head. Alternatively, a receive-only local radio frequency coil can be used in conjunction with body-coil transmission. An RF screen 36 blocks the RF signals from the RF head coil from inducing eddy currents in the gradient coils and the surrounding structures. The resultant radio frequency signals are picked-up by the whole body RF coil 26 , the local RF coil 34 , or other specialized RF coils and demodulated by a receiver 38. With continuing reference to FIGURE 1 and further reference to FIGURE 2, a sequence control circuit 40 controls the gradient pulse amplifiers 20 , the transmitter 24 to generate one of a plurality of pulse sequences loaded from a sequence memory, such as the sequence illustrated in FIGURE 2. A clock or timing generator 42 controls an excitation RF pulse generator 44 and a refocusing RF pulse generator 46. The resonance excitation pulse generator 44 controls the radio frequency transmitter to cause a resonance excitation RF pulse 50 to be applied to the imaging region. The refocusing RF pulse generator 46 causes the radio frequency transmitter to generate a refocusing pulse 52 , e.g., at a time τ after the excitation pulse. The excitation pulse followed by a refocusing pulse causes a spin echo 54 to occur a time 2τ after excitation. In order to create a repeated echo sequence, the refocusing pulse generator 46 further generates one or more additional refocusing pulses 56 at 2τ time intervals after the preceding refocusing pulse. The timing generator 42 also controls a slice select gradient control 60 which controls the gradient amplifiers 20 to cause slice select gradient pulses 62 concurrently with the RF excitation and refocusing pulses. The slice select gradients are used in a single or multiple-slice examination procedure to limit the excitation to a single slice. Alternately, a slab selection pulse may be applied to limit excitation to a slab that is several slices in thickness. The slab is encoded in three-dimensions with a combination of phase and read gradients. The timing control further controls a phase-encode gradient controller 70 and a read gradient control 72 . The phase-encode gradient controller 70 causes the gradient amplifiers 20 to cause an initial phase encode gradient 74 with one of a plurality of phase encode gradient amplitude steps. The read gradient controller causes a first read gradient 76 which is reversed to form an opposite plurality read gradient 78 . The gradient reversal induces first field or gradient echo 80. The read gradient is again reversed to apply a read gradient 82 which is centered about the spin echo 54 . The read gradient is reversed again to an opposite polarity read gradient 84 inducing a second field echo 86 . The phase-encode gradient amplifier applies phase-encode gradients 88, 90 to change the phase-encoding between the field and spin echoes. A phase-unwrapping gradient pulse 92 is equal and opposite to the sum of the phase-encodings applied by phase-encode gradient pulses 74, 88, and 90 such that the net phase-encoding is zeroed prior to application of the net refocusing pulse 56. Thereafter, a like sequence is applied one or more times. In the prior art, the gradient pulse applied during the field echoes 78 and 84 would have been equal in magnitude and duration (but of the opposite polarity) to the read gradient 82 applied during the spin echo as illustrated in phantom in FIGURE 2 at 94 and 96. However, in the present invention, a bandwidth controller 100 adjusts the bandwidth at which each of the field and spin echoes is collected. More specifically, the bandwidth controller 100 controls a read gradient scaler 102 which scales the read gradient profiles, particularly increases its magnitude and shortens its duration relative to the prior art. The bandwidth controller controls a sampling rate controller 104 which controls the sampling rate at which the field echoes are sampled. The sampling rate controller 104 typically adjusts the sampling rate of analog-to-digital converters 106 which digitize the received signal from the field echoes. If the receiver 38 is an analog receiver, the analog-to-digital converters 106 typically follow the receiver. Alternately, the analog-to-digital converters may be disposed between the radio frequency coil and a digital receiver or incorporated within a digital receiver. In the illustrated embodiment, the read gradient scaler 102 doubles the magnitude and halves the width of the profiles of the read gradient pulses 78 and 84 relative to the read gradient pulse 82 applied during the spin echo and the sampling rate controller 104 doubles the sampling rate, e.g., from 32 MHz to 64 MHz. Of course, other proportional adjustments in the read gradient scaling and sampling rates are also contemplated. It is preferred that the area under the field echo and spin echo read gradients remain constant so that the spatial resolution and field-of-view do not change. With continuing reference to FIGURES 1 and 2 and further reference to FIGURE 3, a sorting routine 110 sorts each of the data lines into a data memory 112. More specifically, the sequence control 40 controls the amplitude of phase-encode gradient pulses 74 and 88 relative to each other such that the spin echo has a phase-encoding near the center of k-space, e.g., adjacent a zero phase-encoding between 1/3 of the maximum negative phase-encoding and 1/3 of the maximum positive phase-encoding. The phase-encode gradient 74 and 90 are applied such that the field echoes have phase-encoding between the spin echo phase-encodings and the maximum positive and negative phase-encodings, e.g., between the maximum negative phase-encoding and 1/3 of the maximum negative phase-encoding and between the 1/3 of the maximum positive phase-encoding and the maximum positive phase-encoding. By way of a more concrete example, if there are 256 phase-encoding steps, the phase-encode gradient 74 would be adjusted to cause the first field echo 80 to have a phase-encoding between steps 44 and 128. The second phase-encoding gradient would be adjusted to remove about half of the phase-encoding such that the spin echo 54 has a phase encoding between step -42 and step +43. The phase-encode gradient 88 may be held constant at the equivalent of -85 phase-encoding steps such that a center 1/3 of k-space 114 is filled in the same order as a top 1/3 of k-space 116. Alternately, the second phase-encode gradient 88 may be changed such that the spin echo fills its 1/3 of k-space in a different order. The phase-encode gradient 90 analogously adjusts the phase-encoding such that the second field echo 86 takes on the phase-encodings in a lower 1/3 of k-space 118. Again, the phase-encode gradient 90 may subtract roughly 85 steps of phase-encoding such that the bottom 1/3 is covered in the same order as the upper and center 1/3, or may be varied to cover the bottom 1/3 of k-space in a different order. Of course, the read gradient may be reversed a different number of times to provide only a single field echo or more times to provide more than two field echoes contemporaneously with each spin echo. For example, by doubling the sampling bandwidth and halving the duration of the field echoes, twice as many field echoes can be acquired per spin echo with only the prior art degradation caused by T 2 * phase errors. By increasing the bandwidth still higher, four field echoes can be acquired while still improving the phase-error degradation. When the relative number of field and spin echoes is changed, the segmentation of k-space is changed accordingly. Further, in some applications it may be appropriate to increase the bandwidth of the spin echo sampling and reduce the bandwidth of the field echo samplings. Once the data is sorted into the appropriate k-space format in the data memory 112 , a reconstruction processor 120 reconstructs the data after phase correction using a two-dimensional Fourier transform reconstruction algorithm into an image representation that is stored in an image memory 122 . A video processor 124 retrieves selected portions of the reconstructed image representation and formats it in appropriate format for display on a human-readable monitor 126, such as a video monitor, active matrix monitor, liquid crystal display, or the like. As described above, in the present application the magnetic resonance data lines are acquired with different sampling rates. The final image is reconstructed from a set of data with mixed bandwidth. Although FIGURE 2 describes the invention in terms of a sequence with mixed spin and field echoes, it is to be appreciated that the present invention applies to sequences with one or more field echoes, one or more spin echoes, or combinations thereof. The preferred imaging read-out gradient profile amplitudes G rd (p) during data sampling of an echo (p) are not the same for all echoes. Similarly, the data sampling intervals Δt(p) within each echo are different from echo to echo, depending on the read gradient amplitude. In order for all gradients to view a common field-of-view, the data sampling interval, field-of-view, and the gradient amplitude are related as: G rd (p) FOV Δt(p) = 1 where FOV denotes the field-of-view in the direction of the read-out gradient. A number of complex data points N ds sampled at intervals Δt(p) for each echo is constant. The data sampling time T ds (p) over complete echo (p) is: T ds (p) N ds · Δt(p) where (p) is the echo index. The final two-dimensional complex raw data is obtained by combining such digitally-sampled data into the data memory 112 for later image reconstruction. To minimize discontinuities between segments of k-space, both phase correction and time shift are applied to the data lines belonging to different segments by a correction circuit 130 . For two-dimensional Fourier reconstruction with multiple echoes, the initial imaging excitation by a slice selective RF pulse induces magnetization of a three-dimensional object with a proton nuclear spin density ρ(x,y,z). Assuming that an acquired induction signal or raw data line is labeled by the corresponding view and echo and slice number as: R(t,p,q,s)= ∫ ρ(x,y,z)δ(Z-sΔz)e i(G rd Xt+ΔB(x,y,z)t+ 2πpY FOV )- qTE T 2 dxdy where t,p,q, and s denote the sampling time, view index, echo number, and slice numbers, respectively, G rd (p) is a time dependent magnetic field gradient (read-out) during data acquisition of the echo p, the slice profile denoted by the function δ(z), and ΔB(x,y,z) denotes main field inhomogeneity which includes the contribution of both magnet and susceptibility. The transverse spin relaxation is taken into account by T 2 . Eddy current effect is not included in Equation (3). In order to combine multiple-phase-encoded data lines acquired from different echoes obtained at various echo times into one data area and to form an image free of artifacts, the phase correction on the data is performed by correction circuit 130 . If the data sampling interval is uniform and the read gradient is constant during each data sampling window, data can be directly combined. In practice, each acquired data line is a discrete digital representation of a continuous reality and can be uniform or non-uniform. Due to the rise time constraint of the gradient system, the read-out gradient may have a non-constant shape over part or all of the data sampling window. The acquired data lines may be non-uniform in areas of non-constant read-out gradient. To correct for this non-uniformity, a regridding processor 132 interpolates the non-linear sampled raw data lines to a linear sampled raw data view on a preselected set of uniformly-spaced grid points. Before reconstructing the final image with a two-dimensional Fourier transform, the data from echoes of different bandwidth is properly time-shifted and phase corrected before the last Fourier transform along the phase encoding direction. The noise level of the final composite image is a blend of the noise levels of the relatively noisier high bandwidth sample data and the low bandwidth sample data. More specifically, for a set of data samples obtained as R i , where i=1,2,...,n, with a mean of µ 0 and a variance of σ i 2 , the noise level of the mixed bandwidth data is given by: By placing the lower bandwidth data at the center of k-space, the image is disproportionately weighted in favor of the lower noise data. The present technique can be used with spin echo, field echo, 3D FSE, multiple-shot EPI, GSE, echo-volume imaging, spiral scan and other imaging sequences. For example, as illustrated in FIGURE 4, to speed up a number of field echoes in an echo-planar imaging sequence, the read-out gradient sequence can start with a high amplitude short duration high bandwidth read gradient pulse 140 . In subsequent repetitions 142, 144 ,... the amplitude of the read gradient pulse reduces and its duration increases proportionately such that the field-of-view remains constant. The sampling bandwidth is reduced correspondingly in each field echo, improving the signal-to-noise ratio. Finally, a center read gradient pulse 150 and a corresponding field echo 152 are at the minimum bandwidth. Thereafter, the read-out gradient pulses 154 , 156 ,... start decreasing in duration and increasing in amplitude until a maximum amplitude minimum width read-out gradient pulse 160 occurs contemporaneously with a maximum bandwidth field echo 162 . The phase-encode gradients 170 , 172 ,... are again selected such that the data lines from the lowest bandwidth echoes are disposed near the center of k-space with higher bandwidth data lines being disposed progressively towards the edges. This sequence may be analogously adapted to spin echoes as well.

Claims: A method of magnetic resonance imaging in which magnetic resonance is excited in selected dipoles, and magnetic resonance echoes (54, 80, 86) are induced, the induced magnetic resonance echoes are read-out along a first axis in the presence of a read gradient, the magnetic resonance echoes are digitized and demodulated to form digital data lines, the data lines are stored and reconstructed by a reconstruction algorithm to generate an image representation, which includes the steps of reading-out a first of the magnetic resonance echoes with a first sampling rate in the presence of a first read gradient of a first amplitude to generate a first digital data line with a first bandwidth and reading-out a second of magnetic resonance echoes with a second sampling rate different from the first sampling rate in the presence of a second read gradient of a second amplitude different from the first read-out gradient amplitude to generate a second digital data line of a second bandwidth different from the first bandwidth, the digital data lines of the first and second bandwidths being reconstructed into one resultant image representation, characterised in that the data line with lower bandwidth is phase-encoded with a smaller phase-encoding angle and the data line with the higher bandwidth is phase-encoded with a greater positive or negative phase-encoding angle. A magnetic resonance imaging method as claimed in claim 1, further comprising: following magnetic resonance excitation, applying a first phase-encode gradient along a second axis orthogonal to the first axis; following application of the first phase-encode gradient, reading out the first echo in the presence of the first read gradient to generate the first digital data line; and following the first read gradient, reading out the second echo in the presence of the second read gradient before application of a subsequent excitation of magnetic resonance, whereby the digital data lines of the first and second bandwidth are generated during one excitation of magnetic resonance. A magnetic resonance imaging method as claimed in claim 2, further comprising: after the second read gradient, inducing a third magnetic resonance echo and applying a third read gradient pulse of a third amplitude contemporaneously with the third echo, the third read gradient amplitude being different from the second read gradient amplitude; and reading-out the third echo at a third sampling rate to generate a third data line having a third bandwidth, which third bandwidth is different from said second bandwidth. A magnetic resonance imaging method as claimed in claim 3, wherein the first and third read gradients have a common amplitude greater than the amplitude of the second read gradient and the first and third sampling rates are the same and higher than the second sampling rate such that the first and the third bandwidths are the same and higher than the second bandwidth. A magnetic resonance imaging method as claimed in claims 3 or 4, further comprising: inducing a fourth echo after the third echo and before the subsequent excitation of magnetic resonance, and applying a fourth read gradient of a fourth amplitude in the presence of the fourth echo; and reading-out the fourth echo at a fourth sampling rate to generate a fourth data line having a fourth bandwidth. A magnetic resonance imaging method as claimed in any one of claims 3, 4, and 5, further comprising: exciting magnetic resonance by applying a first radio frequency excitation pulse; applying a radio frequency refocusing pulse; and timing the radio frequency excitation pulse, the radio frequency refocusing pulse, and the first, second, and third read gradients such that the second echo is a spin echo; the first and third echoes being gradient echoes. A magnetic resonance imaging method as claimed in any one of the preceding claims, further comprising: applying phase-encode gradient pulses between adjacent echoes to adjust the phase encoding from echo to echo. A magnetic resonance imaging system which includes a magnet (12) for generating a temporally constant magnetic field through an examination region (14), gradient magnetic field coils (22) and a gradient magnetic field controller (70, 72) for generating at least phase and read magnetic field gradient pulses in orthogonal directions across the examination region, a radio frequency pulse controller (44, 46) and transmitter (24) for inducing dipoles in the examination region to resonance such that radio frequency resonance signals are generated, a receiver (38) for receiving and demodulating the radio frequency magnetic resonance signals to produce a series of data lines, the receiver being associated with an analog-to-digital converter (106) such that the data lines are digital, a data memory (112) for storing the digital data lines, and a reconstruction processor (120) for reconstructing the data lines from the data memory into an image representation, including a read gradient scaler (102) for scaling amplitude and width of selected read gradient pulses such that a bandwidth of the radio frequency magnetic resonance signals received during the scaled read gradient is altered; and a sampling rate control (104) for controlling a sampling rate of each received radio frequency magnetic resonance signal in accordance with the change in bandwidth, wherein the gradient magnetic field and radio frequency pulse controllers are controlled such that data lines with the lowest bandwidth are phase-encoded with the smallest phase-encoding angles and data lines with the greatest positive and negative phase-encoding angles have the highest bandwidth. A magnetic resonance imaging system as claimed in claim 8, further comprising: a data line sorter (110) for sorting the digital data lines in accordance with bandwidth, the sorted data lines being supplied to the data memory. A magnetic resonance imaging system as claimed in claim 8 or claim 9, wherein the first and a third read gradients have a common amplitude greater than the amplitude of the second read gradient, and the first and a third sampling rates are the same and higher than the second sampling rate such that the first and a third bandwidths are the same and higher than the second bandwidth.

Inventors: GULLAPALLI RAO P, LIU HAIYING, LONCAR MARK J, GULLAPALLI, RAO P., LIU, HAIYING, LONCAR, MARK J.
Applicants: MARCONI MEDICAL SYS INC, MARCONI MEDICAL SYSTEMS, INC.
Application year: 1996
Main classification: G01R33
IPC codes: A61B5, G01R33