Electricity: measuring and testing – Particle precession resonance – Using a nuclear resonance spectrometer system
Reexamination Certificate
2001-04-04
2002-06-11
Arana, Louis (Department: 2862)
Electricity: measuring and testing
Particle precession resonance
Using a nuclear resonance spectrometer system
C324S307000
Reexamination Certificate
active
06404196
ABSTRACT:
BACKGROUND OF THE INVENTION
The present invention relates to magnetic resonance (“MR”) imaging. It finds particular application in conjunction with correcting MRI motion artifacts and main field fluctuation and will be described with particular reference thereto. It will be appreciated, however, that the invention is also amenable to other like applications.
Magnetic resonance imaging is a diagnostic imaging modality that does not rely on ionizing radiation. Instead, it uses strong (ideally) static magnetic fields, radio-frequency (“RF”) pulses of energy and magnetic field gradient waveforms. More specifically, MR imaging is a non-invasive procedure that uses nuclear magnetization and radio waves for producing internal pictures of a subject. Three-dimensional diagnostic image data is acquired for respective “slices” of an area of the subject under investigation. These slices of data typically provide structural detail having a resolution of one (1) millimeter or better.
Programmed steps for collecting data, which is used to generate the slices of the diagnostic image, are known as an MR image pulse sequence. The MR image pulse sequence includes magnetic field gradient waveforms, applied along three (3) axes, and one (1) or more RF pulses of energy. The set of gradient waveforms and RF pulses are repeated a number of times to collect sufficient data to reconstruct the slices of the image.
The data for each slice is acquired during respective excitations of the MR device. Ideally, there is little or no variations in the phase of the nuclear magnetization during the respective excitations. However, movement of the subject (caused, for example, by breathing, cardiac pulsation, blood pulsation, and/or voluntary movement) and/or fluctuations of the main magnetic field strength may change the nuclear magnetization phase from one excitation to the next. This change in the phase of the nuclear magnetization may degrade the quality of the MR data used to produce the images.
A non-phase encoded additional echo signal, prior to or after the data echo used for image generation, may be used to detect view dependent global phase variations when two-dimensional Fourier transform encoding and reconstruction algorithms are used. This “Navigator” echo passes through the center of the data space (K-space) each time, while the MR image data is ordered sequentially and linearly. Then, computational methods are used to correct the undesired view-to-view phase variation, thereby eliminating a significant source of image artifacts.
With reference to
FIG. 1
, a typical MR imaging pulse
10
includes a slice select (frequency encoding) gradient
12
and an RF pulse
14
(i.e., the actual MR image signal). The slice select gradient
12
and the RF pulse
14
define a spatial location in which the image data occurs. A phase gradient
16
and a read gradient
18
determine how data is acquired in K-space, which is used to relate the raw data to the final image. The time interval between successive pulse cycles (“TR”) is indicated as interval
19
. The time interval from one pulse to the measurement of the MR signal is indicated as interval
20
. The MR data is acquired during the interval
21
.
The most common method for spatial encoding and reconstructing an MR image is called an N-dimensional Fourier transform (“N-DFT”). Two-dimensional Fourier transforms (“2-DFT”) are used more than 90% of the time while three-dimensional Fourier transforms (“3-DFT”) are used less than 10% of the time. On very rare occasions, a completely different strategy for encoding and reconstructing the data is used (e.g., wavelets, singular value decomposition, etc). Although, for the sake of simplicity, the MR imaging procedure is only described in terms of a 2-DFT comparison, it is to be understood that Fourier transforms of other dimensions are also contemplated.
In a 2-DFT MR imaging procedure, the MR image data is collected in a checkerboard fashion. Rows of the MR image data space are selected according to the phase gradient
16
. The frequency encoding gradient
12
which, when combined with the actual MR image signal
14
, defines the columns in the MR image data space.
FIG. 2
illustrates a conventional K-space
22
, including a Kx axis
24
and a Ky axis
26
, corresponding to the sequence
10
shown in
FIG. 1. A
K-space center
28
, which defines an average signal amplitude in the image, is located at the intersection of the Kx and Ky axes
24
,
26
, respectively. A line
32
illustrates the path traversed by the typical 2-DFT MR image gradients used in the data acquisition sequence illustrated in
FIG. 1. A
new row
32
a
is collected following each repetition time (“TR”) and an image is not usually reconstructed via 2-DFT until the entire data space is filled. Each point in the K-space is a complex number and, therefore, includes both real and imaginary (magnitude and phase) parts.
There are many factors that can alter the phase of the MR image signal. For example, motion through the MR image gradient waveforms and/or variations in the strength of the magnetic field during the total time of the MR image acquisition.
Because it defines an average characteristic of the object being imaged, the center of the K-space
22
is important in MR imaging. Ideally, over the course of the examination, the anatomy of the subject does not change significantly. If this were the case, no point would vary in magnitude or phase even if it is acquired multiple times. Because the center point defines the greatest amplitude in the raw data space, changes in the tissue being imaged, which may occur during TR, manifest themselves as variations in this value. Unfortunately, this point is only collected once during each normal 2-DFT acquisition. U.S. Pat. No. 4,937,526 discloses a method for forming an MR image signal prior to or after the imaging data acquisition. The trajectory in the K-space and the gradient waveforms for such an acquisition are shown in
FIGS. 3 and 4
.
FIG. 3
illustrates a typical MR imaging pulse with a Navigator echo
50
. The pulse
50
includes a slice select (frequency encoding) gradient
52
and an RF pulse
54
(i.e., the actual MR image signal). The slice select gradient
52
and the RF pulse
54
define a spatial location in which the image data occurs. A phase gradient
56
and a read gradient
58
determine how data is acquired in K-space, which is used to relate the raw data to the final image. The time interval between successive pulse cycles (“TR”) is indicated as interval
59
. The time interval from one pulse to the measurement of the MR signal is indicated as interval
60
.
The first portion of the TR cycle
59
shown in
FIG. 3
appears to be a conventional 2-DFT acquisition. However, Navigator pulses
62
cause the supplementary trajectory to cross the center of K-space in each of the TR cycles
59
. Because the time interval between the imaging and Navigator echo is short, it is assumed that no other source of phase variation occurs beyond that imposed on the imaging data.
After the phase of each Navigator echo (i.e., the center of K-space)
62
is determined, an average is calculated across all views. The difference between the mean phase over all views and the phase of each specific view is determined and removed by appropriate well known mathematical methods. The corrected MR image data is then reconstructed to form a reduced artifact image. Again, the additional gradient waveforms are required to take the acquisition trajectory through the center of K-space so that phase information at the center of K-space can be determined for each TR. These additional gradients add time to each acquisition and reduce the temporal efficiency of the scanning while compensating for motion artifacts. This loss of efficiency is particularly detrimental to short TR rapid scanning methods routinely used in MR imaging. The loss of efficiency is often too severe to warrant their use in these common imaging methods.
FIG. 4
illustrates a K-space
64
, including a Kx axis
66
and a Ky axis
68
, corresponding to
Chung Yiu-Cho
Duerk Jeffrey L.
Lewin Jonathan S.
Merkle Elmar
Shankaranaravanan Ajit
Arana Louis
Case Western Reserve University
Fay Sharpe Fagan Minnich & McKee LLP
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