Electricity: measuring and testing – Particle precession resonance – Determine fluid flow rate
Reexamination Certificate
2002-12-06
2004-08-17
Arana, Louis M. (Department: 2859)
Electricity: measuring and testing
Particle precession resonance
Determine fluid flow rate
C324S309000, C600S410000
Reexamination Certificate
active
06777933
ABSTRACT:
This invention relates to a method of reducing the effects of motion of objects in an image, in particular to a method of reducing the effects of motion in magnetic resonance imaging. The invention particularly relates to a method of compensating for patient motion to produce a focussed image.
Magnetic resonance imaging or MRI is a well known medical imaging technique. In essence the technique relies on the reaction of the magnetic moments of certain nuclei to applied magnetic fields. Protons and neutrons, the basic constituents of nuclei, posses magnetic dipole moments. In nuclei with an even number of protons and an even number of neutrons the net effect is no residual magnetic moment. However nuclei with uneven atomic number (or uneven atomic mass) have a net magnetic dipole and hence a magnetic moment. At room temperature in the absence of an external magnetic field one would expect to find a random orientation of magnetic moments in a medium.
In an MRI imaging system an intense magnetic field is applied to the area to be imaged. This field is applied in one direction, conventionally referred to as the z-direction. The effect of the applied field is to align the magnetic dipoles in the item being imaged. The dipoles do not all line up in exactly the same way however. The dipoles tend to adopt either an orientation lined up in the same direction as the field, referred to as parallel, or an orientation where the dipoles align opposite the field direction, the antiparallel orientation. At room temperature, due to the parallel state being slightly more energetically favourable, slightly more nuclei tend to adopt the parallel configuration than the antiparallel configuration. This results in a net overall magnetic moment for the medium, parallel to the applied field.
The coupling effects of the magnetic moment of the nuclei with the applied field does not cause an exact alignment of the nuclear moment with the applied field. Instead the magnetic moment processes around the applied field. The frequency of precession, called the Larmor frequency, is proportional to the strength of the applied field. The stronger the applied field the faster the rate of precession.
In effect one can consider that the dipole moments of the nuclei have aligned so there is a component of the moment in the z-direction and a component rotating in the x-y plane at the Larmor frequency. As mentioned, throughout the whole object being imaged there is a greater component parallel to the z-direction than antiparallel so there is a net moment for the object. However the components in the x-y plane are still randomly arranged in the presence of a single field so there is no net moment in the x-y plane.
Applying an RF magnetic field at the Larmor frequency perpendicular to the applied field causes the dipoles to tip into the transverse, or x-y, plane. It also causes alignment of the dipoles. The net result is then a net magnetic moment in the x-y plane rotating at the Larmor frequency.
When the RF field is removed this net magnetic moment can be measured due to the inductance caused in receiver coils. Of course once the RF field is removed the net magnetisation of the item being imaged will start to revert to what it had been as the magnetic moments of the nuclei begin to align with the z-direction again.
There are two separate decay processes that occur. The first is the increase in the z-direction component of overall magnetic moment. Thus is sometimes referred to as longitudinal or spin axis relaxation and is due to the transfer of energy between excited nuclei and the lattice, or nearby macro-molecules. The second process, which is independent of the first, is that the precession of the moments of the nuclei, which had been brought into in phase by the transverse rf field, start to de-phase reducing the x-y component. The de-phasing process. known as transverse relaxation or spin-spin interaction is due to transfer of energy between nuclei in different states and also from magnetic field inhomogenities. In both decay processes the different types of material present in an object, say the differing types of tissue in a patient, will affect the relaxation processes. Hence measuring the differing strengths of signal received by the different types of tissue will reveal contrast in an image.
In order to form an image it is necessary to encode the dipoles of the signals emitted by the nuclei after magnetization have information regarding to the spatial positioning of those nuclei. The imaging processes can usually be described in the following terms. First of all is the step of selecting an image slice, i.e. a small volume to be imaged, and then spatially encoding the magnetic resonance signal emanating from that slice. The basis for this is that the frequency at which a nucleus resonates, its Larmor frequency, is a function of the strength of the static magnetic field in which it is located. Therefore by altering the strength of the magnetic field as a function of position, i.e introducing a magnetic field gradient, the Larmor frequency will also vary as a function of position.
Typically therefore a weak magnetic field that changes linearly with position is superimposed on the main static field to create a magnetic field gradient along the z-direction. An RF pulse with a narrow range of frequencies is then applied transversely. Only those nuclei whose Larmor frequency matches the frequency of the applied RF pulse will actually absorb the RF energy and undergo the tipping and alignment described above. Therefore by a careful choice of RF frequency only a narrow band or slice of the object being imaged will be excited.
Having selectively excited a slice of the object to imaged it is necessary to achieve spatial resolution within in a slice. Spatial resolution in one dimension, say the x-direction, can be achieved through use of a frequency encoding gradient. Immediately following the RF excitation pulse all spins of the nuclei of interest within the selected slice will be processing at the same frequency. Application of an additional gradient, orthogonal to the z-direction gives spatial resolution in one dimension. This additional gradient, known as a frequency encode gradient, will alter the Larmor frequency of the spin precession across the slice and allow spatial resolution.
Note that for medical MRI the nuclei of interest is almost exclusively the nucleus of hydrogen. However other nuclei species could be of interest in certain applications.
To get two dimension resolution across the slice it is necessary to use a phase encode step as well. Here following the RF excitation pulse a phase encoding gradient is applied in the y-direction for a short time. Remember that immediately following the RF excitation pulse all the spins in the selected slice will be in phase and processing at the same frequency. If a phase encode gradient is applied in the y-direction the spins will have their resonant frequencies, and hence the rate of procession, altered according to their position along the y-direction. When the phase enode gradient is removed all nuclei in the slice will again be subject to the same static field strength and hence the spins will again start to precess at the same frequency. The effect of the phase encode gradient will have been to alter the phase of the spins according to their position along the y-axis in a known manner. The frequency encode gradient may then be re-applied.
The measured signal at a particular frequency (and therefore position along the x-axis) is the sum of all the vector contributions from a row of spins in the y-direction. The actual signal measured of course is a composite of all the frequency components along the x-axis.
To generate an image during the time that the frequency encode gradient is applied the signal is sampled N
x
times yielding a pe-line, which is a vector or line of data having N
x
points. Repeating the measurements N
y
times for differing values of the y-gradient yields a matrix of N
x
×N
y
amplitude points. In general to generate a final image of N&t
Arana Louis M.
Nixon & Vanderhye P.C.
QinetiQ Limited
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