Patent Application
20060084861
This application is related to a patent application titled “Expanding MRI Probe,” attorney's docket number 334/03529, filed on even date, at the US Patent and Trademark Office, the disclosure of which is incorporated herein by reference.
The field of the invention is nuclear magnetic resonance probes.
Problems with Conventional MRI
Conventional MRI (magnetic resonance imaging) systems suffer from a number of limitations. They require highly homogeneous magnetic fields, which, for imaging a large volume such as the human body, generally means large and expensive equipment that is not very mobile. The distance between the imaged volume and the RF antenna means that a rather long acquisition time is needed to obtain a reasonable signal to noise ratio at high resolution. Because most patients cannot tolerate being inside the narrow bore of a large magnet for more than a few minutes, the images have limited resolution, typically about 1 millimeter. While “open” MRI magnets exist, with less claustrophobic bores, they have lower magnetic fields which further reduces the signal to noise ratio, so these systems also typically have resolution no better than 1 mm. Conventional MRI systems are thus unable to resolve plaque in the thin walls of arteries, even though MRI, in contrast to other imaging techniques, is very good at distinguishing between different types of soft tissues.
“Inside Out” MRI
These limitations of conventional MRI have led to the development of newer MRI techniques, in which smaller volumes can be imaged with higher resolution, and/or with less expensive and more portable equipment. Often, the region being imaged is outside the magnet and the RF coil, rather than being surrounded by the magnet and RF coil, as in conventional MRI.
One problem in “inside out” medical MRI, is that a high magnetic field gradient is generally produced outside the magnet. For a fixed small RF bandwidth typical of those used in conventional MRI, higher field gradients make the resonant region narrower, with fewer nuclei to produce a signal, and allow the nuclei to diffuse away from the resonant region quickly, further reducing the signal. If the bandwidth is made wider, then the noise increases. In either case, the signal to noise ratio (SNR) is small. This problem can also be at least partly overcome by using appropriate magnet configurations. One such configuration is described by Jackson, U.S. Pat. No. 4,350,955, the disclosure of which is incorporated herein by reference. Jackson describes two magnets, arranged along the z-axis with a gap between them, and magnetized in opposite directions parallel to the z-axis. This configuration produces a ring surrounding the probe where the magnetic field intensity has a saddle point, at which the magnetic field gradient is relatively small. Other magnet configurations which produce a saddle point in the magnetic field intensity outside the probes are described by Clow, U.S. Pat. No. 4,629,986, by Masi, U.S. Pat. No. 4,717,876 (including both axially and radially magnetized magnets in an axisymmetric probe), by Locatelli U.S. Pat. No. 5,610,522, and by Kleinberg et al, “Novel NMR Apparatus for Investigating an External Sample”, J. Magn. Res., 97, p. 466, 1992, the disclosures of which are incorporated herein by reference.
Medical MRI Receiver Probes
Most MRI probes used for medical applications are not self-contained probes with a magnet and RF transmitter, but only have an RF receiver, which is used inside the body to pick up signals from a small region of interest, while a conventional MRI magnet and RF transmitter are located outside the body to excite the region. Such probes are described by Atalar, U.S. Pat. No. 5,699,801, by Bradley, U.S. Pat. No. 5,050,607, by Kandarpa et al., J. Vasc. and Interventional Radiology, 4, pp. 419-427, 1993, and by H. H. Quick et al, Magnetic Resonance in Medicine 42:738-745, 1999, the disclosures of which are incorporated herein by reference. Sometimes the probe also produces a local field gradient, either using coils, described by Young, U.S. Pat. No. 5,303,707, or using pieces of soft magnetic material, described by Golan, U.S. Pat. No. 6,377,048, the disclosures of which are also incorporated herein by reference.
Self-Contained Medical MRI Probes
Pulyer, U.S. Pat. No. 5,572,132, the disclosure of which is incorporated herein by reference, describes a medical MRI probe which is self-contained, including a magnet, gradient coils, and RF transmitting and receiving coils. It has magnets (two magnets both magnetized in the same direction along the z-axis, with a gap between them) carefully shaped to produce a limited region of low magnetic field gradient outside the probe. The probe can be used for NMR spectroscopy, as well as for imaging.
Throughout this application, we will refer to the longitudinal direction, for example in a blood vessel, as the z direction, with the x and y directions perpendicular to the z direction and to each other.
Westphal et al, in U.S. Pat. No. 5,959,454, the disclosure of which is incorporated herein by reference, describes an MRI probe with an external imaging region on one side, for use outside the body to examine skin, for example.
Cho, U.S. Pat. No. 5,023,554, and Kikinis, U.S. Pat. No. 5,390,673, the disclosures of which are incorporated herein by reference, describe using a very inhomogeneous static magnetic field, and imaging slices that are far from flat, for medical MRI.
Crowley, U.S. Pat. No. 5,304,930, the disclosure of which is incorporated herein by reference, describes an MRI device located just outside the body, and used to image a region of the body. Prado et al, in U.S. Pat. No. 6,489,767, the disclosure of which is incorporated herein by reference, describes a palm-sized MRI probe with a planar imaging region on one side.
The NMR-MOUSE, a mobile NMR sensor, is described by Todica et al, J. Magn. Res. 164 (2003) 220-227, by Klein et al, J. Magn. Res. 164 (2003) 310-320, and by references therein. Anferova et al, “Construction of a NMR-MOUSE with Short Dead Time,” Concepts in Magnetic Resonance (Magnetic Resonance Engineering) 15(1), 15-25 (2002) describes ways of designing the coil and other components in an NMR-MOUSE which result in a dead time in the RF receiver amplifier of only 20 microseconds, after transmitting RF pulses through the same antenna. The RF frequency is about 20 MHz. This dead time is much shorter than the dead times in conventional MRI systems which use the same RF antenna for transmission and receiving, as described, for example, by Eiichi Fukushima and Stephen B. W. Roeder, in Experimental Pulse NMR: A Nuts and Bolts Approach, Perseus Publishing, 1986. The disclosures of all these articles and this book are incorporated herein by reference.
In U.S. Pat. No. 6,704,594, the disclosure of which is incorporated herein by reference, Blank et al describe a self-contained intravascular MRI probe. The probe uses two cylindrical magnets, arranged along the z-axis, magnetized in opposite directions perpendicular to the z-axis. This configuration, with an RF transmitting and receiving coil on one side of the probe, produces sector-shaped imaging slices on that side of the probe, with limited axial and radial extent.
Non-axisymmetric Magnet Configurations
An aspect of some embodiments of the invention concerns self-contained MRI probes, for example intravascular probes, which have novel magnet and/or RF coil configurations that may result in improved properties compared to the prior art. Improved properties may include one or more of: higher static magnetic field in the imaging region for a given strength magnet; greater radial penetration of the static magnetic field and the RF field into the region surrounding the probe; a field of view with greater axial extent; and the ability to image regions in more than one azimuthal direction, and/or at one more than one axial position, simultaneously, without any need to rotate the probe or to move the probe axially.
In some of these novel configurations, there are a plurality of magnets arranged along the z-axis of the probe, configured so that the magnetic field is not axisymmetric, and adjacent magnets have directions of magnetization that differ by an angle that is substantially different from 0 degrees or 180 degrees. For example, the directions differ by an angle that is between 10 degrees and 170 degrees. Optionally, the directions differ by an angle that is between 20 degrees and 160 degrees, or between 40 degrees and 140 degrees.
For example, there are two magnets, the bottom one being magnetized at an angle of 45 degrees between the +z direction and the +x direction, and the top one being magnetized at an angle of 45 degrees between the +z direction and the −x direction. This configuration, in which the direction of magnetization differs by 90 degrees for the two magnets, produces a magnetic field just to the +x side of the probe that is greater than the magnetic field just to the −x side of the probe, and greater than the field would be just outside both sides of the probe if the two magnets were magnetized in opposite directions along the x-axis.
In another example, there are a plurality of magnets arranged axially along the probe. Each magnet is magnetized in a direction perpendicular to the z-axis, but in a different direction in the x-y plane that differs by less than 180 degrees from the direction of magnetization of the next magnet. For example, there are four magnets, magnetized respectively in the +x direction, the +y direction, the −x direction, and the −y direction. Each of these magnets has its own RF coil, and is used to obtain imaging data from a different azimuthal direction. Since all four magnets can produce data simultaneously, there is no need to take data first in one azimuthal direction, then rotate the probe, then take data in another azimuthal direction, etc., and a complete scan can be done more quickly. Even if the data from all the different magnets is not obtained simultaneously, but some or all of it is obtained sequentially, for example to avoid interference between magnets, this probe configuration still saves time because it is not necessary to rotate the probe. It also provides greater accuracy, because the difference in azimuthal direction between the different magnets is fixed by the structure of the probe, while in a probe which must be rotated, there may be an error in the angle of rotation.
Although the different magnets are located at different axial positions, this may not make much difference in certain applications, for example in examining arteries for plaque, because plaque tends to extend some distance along arteries at a same azimuthal location. Optionally, additional magnets are located further away axially, where the plaque is likely to look different, and the additional magnets are used to obtain data at different axial locations, simultaneously or sequentially, without the need to move the probe axially.
Optionally, in these probes, there is a structure which allows the probe to expand radially, for example inside a blood vessel, pressing each magnet, and/or its associated RF coil, against the wall of the blood vessel in a different azimuthal direction. Optionally, each magnet presses against the wall with its direction of magnetization normal to the wall at that point, maximizing the field intensity in the imaging region, for each position around the wall azimuthally. Alternatively, each magnet presses against the wall with its direction of magnetization parallel to the wall, and with an associated RF coil adjacent to the wall, producing an RF magnetic field normal to the wall.
Configurations with Longitudinally Magnetized Magnets
An aspect of some embodiments of the invention concerns self-contained MRI probes, for example intravascular probes, which have two magnets arranged along the z-axis, with direction of magnetization respectively in the +z and −z direction, and substantially no gap between the magnets. With no gap between the magnets, there is no saddle point in the magnetic field, surrounded by a region of low field gradient, within the field of view of the probe. Alternatively, there is a gap between at least some adjacent magnets, but the region of low field gradient is still outside the field of view.
Alternatively, there are three, four, or more magnets arranged along the z-axis, with direction of magnetization alternately in the +z and −z direction. These configurations, extended with a sufficiently large number of magnets magnetized alternately in the +z and −z directions, make it possible to produce a field of view which extends an arbitrary distance longitudinally. Optionally, there is a region of low field gradient, even within the field of view.
In some embodiments of the invention, various coil configurations are used, where optimal configurations make efficient use of the static magnetic field and RF magnetic field to obtain NMR data. Optimally, to make efficient use of the fields, the RF magnetic field should be close to perpendicular to the static magnetic field, and close to its maximum intensity, over most of the volume where the static magnetic field is comparable to its maximum intensity. Optionally, this volume of maximum or near maximum fields extends over a length that is a significant fraction of the radial dimension of the probe, radially out to a distance beyond the surface of the probe that is at least a significant fraction of the radius of the probe, and over a significant range of azimuthal angles. Efficient use of the static and RF magnetic fields increases the SNR, or reduces the RF heating of tissue for each bit of imaging data at a given SNR, and increases the field of view for a given probe size.
Other Magnet Configurations
An aspect of some embodiments of the invention concerns a magnet for a self-contained MRI probe. The magnet is magnetized in the x-direction, and has a circular cross-section but excluding one or two sub-volumes on their periphery. The sub-volumes may, for example, be volumes outside planes that are, for example, delineated by chords, parallel to the x-axis, on one or both sides of the circumference. In this case, the missing pieces of the magnet would contribute very little to the magnetic field strength just outside the magnet near the x-axis, which is the location where the field is greatest, and where the imaging region is located if the magnet is being used efficiently. Optionally, the volume that these missing sub-volumes took up is used for electronic circuitry, or for a balloon. Alternatively, the imaging region is located outside the magnet near the y-axis, but there is only one missing sub-volume, on the side of the magnet opposite the imaging region. In this case too, the missing sub-volume of the magnet would contribute relatively little to the magnetic field in the imaging region.
Optionally, the circular cross-section of the magnet is not a solid circle but a hollow circle, and, with the missing piece removed, the cross-section of the magnet is C-shaped. The hollow in the center is used, for example, to allow blood to flow through the probe, for a guide wire, and/or for cables which connect to the RF antenna.
An aspect of some embodiments of the invention concerns an MRI probe with a long cylindrical magnet (not necessarily a circular cylinder), and with an end cap at one or both ends. The end cap is optionally made of a high permeability material such as iron, and/or is sufficiently thick, and has high enough saturation flux density, to carry a substantial fraction of the flux of the magnet within one diameter of the end. The end cap may make the magnetic field around the magnet more uniform as a function of longitudinal position, possibly over most of the length of the magnet, and may make the field fall off more abruptly near the end of the magnet. This may make the imaging region, which may be limited by the contours of field strength, more uniform over the length of the magnet, potentially improving the signal to noise ratio. A more uniform imaging region may also provide more accurate radial voxel assignment for imagining the blood vessel wall, therefore making it possible to accurately estimate the distance of the plaque from the edge of the lumen which is an important parameter in evaluating its vulnerability. The magnet is magnetized substantially perpendicular to the axis of the cylinder.
An aspect of some embodiments of the invention concerns an MRI probe with a cylindrical permanent magnet and an RF coil, with the RF coil located in a shallow depression in the surface of the magnet, rather than located outside the outer diameter of the magnet. The parts of the magnet which extend to the same outer diameter as the RF coil, to the sides of the RF coil, are referred to herein as “ears.” Depending on the dimensions of the coil and the magnet, this configuration produces a higher static magnetic field in the imaging region of the probe, and hence higher resolution, higher signal to noise ratio, or shorter acquisition time, for the same permanent magnet material and the same outer envelope of the probe.
Reduced Eddy Currents
An aspect of an embodiment of the invention concerns an MRI probe with reduced eddy currents. The probe comprises an RF coil (or another kind of RF antenna or RF transmitting element) and a permanent magnet, with a design which reduces eddy currents induced in the magnet at the RF frequency when the RF coil transmits RF power. In an exemplary embodiment of the invention, there is a gap between the RF coil and the magnet. In an exemplary embodiment of the invention, there is a layer of a good conductor, such as copper, between the RF coil and the magnet, which shields the magnet, but not the imaging volume, from the RF fields generated by the RF coil. Although the conductor will also have eddy currents, they will generally dissipate less power than eddy currents in the lower conductivity magnet without shielding, since the skin depth is generally small. In an exemplary embodiment of the invention, the magnet is laminated, or has slots in it. Optionally, any gap between the RF coil and the magnet, or any layer of conductor between the RF coil and the magnet, is thick enough to significantly reduce eddy currents in the magnet. But optionally, the gap is not so thick as to significantly reduce the static magnet field in the imaging region by reducing the volume of the magnet for a given probe envelope, or by increasing the distance between the magnet and the imaging region. Optionally, the gap and/or the thickness of the conductor is optimized by minimizing a function that reflects both the adverse effect of eddy currents in the magnet, and the adverse effect of reducing the static magnetic field in the imaging region.
Non-imaging NMR with Self-contained Probes
Any of the MRI probes described above are also optionally used for non-imaging NMR. For example, instead of creating an image by obtaining NMR data from different longitudinal, azimuthal or radial positions relative to the probe, optionally the NMR data from all positions is lumped together, to obtain information about an average, possibly a weighted average, of the NMR characteristics of material in the field of view of the probe. This is done, for example, in order to increase the SNR of the signal, or to decrease the acquisition time for a given SNR, at the cost of losing information about the spatial distribution of the material. Losing information about the spatial distribution might not matter very much if it is expected that the material is distributed fairly broadly over the field of view of the probe. The NMR characteristics comprise, for example, the density of protons and/or other nuclei, T1, T2, the diffusion rate, and/or spectroscopic data. Such characteristics are optionally used, for example, to distinguish plaque from healthy tissue in the walls of blood vessels, even without imaging.
Such non-imaging NMR data may be obtained with any self-contained NMR or MRI probe, not just with the probes described above, including probes for which the static magnetic field has a saddle point around which the field is locally uniform. Although the presence of a locally uniform magnetic field region may make the probe particularly suitable for obtaining spectroscopic data, such a probe is also useful for obtaining data on density, T1, T2, and diffusion rate. An aspect of an embodiment of the invention concerns an NMR system using a self-contained NMR probe, whether the probe is suitable for imaging or not, which system is used to measure spatially averaged non-imaging NMR characteristics, other than spectroscopic data, and where there is a saddle point in the static magnetic field outside the probe.
RF Antenna with Short Dead Time
An aspect of some embodiments of the invention concerns an active protection circuit with very high bandwidth. The protection circuit isolates a sensitive low noise amplifies used to amplify weak received NMR signals, from an RF antenna, when the antenna is transmitting high power RF pulses. The same RF antenna can thus be used for both receiving and transmitting. The protection circuit uses active elements, such as toroid protectors, and optionally also uses passive elements such as a high pass filter and Schottky diodes. The circuit has a very high bandwidth and short ringing time, allowing the amplifier to have a very short dead time, as short as a few microseconds.
There is thus provided, in accordance with an exemplary embodiment of the invention, a probe, with a longitudinal axis, for use in an NMR system, the probe comprising:
Optionally, said adjacent static magnetic field sources are displaced from each other along the longitudinal axis.
Optionally, adjacent static magnetic field sources are magnetized in directions that differ by more than 20 degrees and less than 160 degrees.
Optionally, adjacent static magnetic field sources are magnetized in directions that differ by more than 40 degrees and less than 140 degrees.
In an embodiment of the invention, the probe is adapted for inserting into a cavity in the body.
Optionally, the probe is adapted for inserting into a blood vessel.
Optionally, the probe is adapted for inserting into a blood vessel with inner diameter between 1.5 mm and 6 mm.
Optionally, the probe is adapted for inserting into a blood vessel with inner diameter between 2 mm and 4 mm.
In an embodiment of the invention, the static magnetic field sources comprise a first magnetic field source and a second magnetic field source, both with longitudinal components of magnetization having a same sign, and with transverse components of magnetization differing in direction by more than 90 degrees.
Optionally, there is a gap between the first and second magnetic field sources.
Optionally, the transverse components of magnetization differ in direction by more than 140 degrees.
Optionally, the transverse components of magnetization differ in direction by more than 160 degrees.
Optionally, the ratio of the magnitude of the transverse and longitudinal components of magnetization is greater than 0.5 and less than 2, for both the first and second magnetic field sources.
Optionally, the ratio is between 0.8 and 1.2, for both the first and second magnetic field sources.
In an embodiment of the invention, at least one of the at least one antennas extends over a range in the longitudinal direction that overlaps the longitudinal ranges of both the first and second magnetic field sources, and is located on one side of the longitudinal axis.
Optionally, the center of said antenna is located within 60 degrees of the location at which the longitudinal component of the static magnetic field is greatest, for that longitudinal position and distance from the longitudinal axis.
Optionally, the center of said antenna is located within 30 degrees azimuthally of said location.
In an embodiment of the invention, the first and second magnetic field sources extend radially to the surface of a smallest convex volume which includes both magnetic field sources, except for a slot carved into one or both of the first and second magnetic field source, and said antenna is located in one or both slots, entirely within said smallest convex volume.
Optionally, the smallest convex volume is cylindrical.
Optionally, the static magnetic field sources each have a component of magnetization transverse to the longitudinal axis that has a magnitude more than 2 times the magnitude of the longitudinal component of magnetization.
Optionally, the transverse component has a magnitude more than 5 times the magnitude of the longitudinal component.
Optionally, the transverse components of magnetization of adjacent static magnetic field sources differ in direction by more than 40 degrees and less than 140 degrees.
Optionally, the at least one antennas comprise an antenna associated with each of the static magnetic field sources.
Optionally, for each of said antennas, the static magnetic field in the extended sub-region is at least 80% produced by the static magnetic field source which that antenna is associated with.
Optionally, each sub-region has a limited range of azimuthal angles, and the azimuthal direction of the center of the range differs by more than 40 degrees and less than 140 degrees for at least two antennas associated with adjacent static magnetic field sources.
Optionally, the azimuthal direction of the center of the range for each of said antennas differs from the transverse component of the direction of magnetization (or the direction opposite to the direction of magnetization) of the static magnetic field source associated with that antenna by a same angle, to within ±20 degrees.
Optionally, the azimuthal direction of the center of the range for each of said antennas differs from the transverse component of the direction of magnetization (or the direction opposite to the direction of magnetization) of the static magnetic field source associated with that antenna by less than 20 degrees.
Alternatively, the azimuthal direction of the center of the range for each of said antennas differs from the transverse component of the direction of magnetization (or the direction opposite to the direction of magnetization) of the static magnetic field source associated with that antenna by between 70 and 110 degrees.
Optionally, the azimuthal direction of the center of the range for each of said antennas differs from the direction of the transverse component of the time-varying magnetic field produced by that antenna in the center of its sub-region, by less than 20 degrees.
Alternatively, the azimuthal direction of the center of the range for each of said antennas differs from the direction of the transverse component of the time-varying magnetic field produced by that antenna in the center of its sub-region, by between 70 and 110 degrees.
In an embodiment of the invention, the set of all azimuthal directions that are included within the range of any of said antennas does not have a gap greater than 90 degrees.
Optionally, the set of all azimuthal directions that are included within the range of any of said antennas does not have a gap greater than 45 degrees.
Optionally, the set of all azimuthal directions that are included within the range of any of said antennas covers more than 180 degrees.
Optionally, the set of all azimuthal directions that are included within the range of any of said antennas covers 360 degrees.
In an embodiment of the invention, the probe includes an expansion mechanism which, when it expands, moves at least two of the magnetic field sources, and its associated antenna, in different directions transverse to the longitudinal axis.
Optionally, the expansion mechanism moves each of the at least two static magnetic field sources in a direction that differs from the azimuthal direction of the center of the range for the antenna which that static magnetic field source is associated with, by a same angle, to within ±20 degrees.
Optionally, the expansion mechanism moves each of the at least two static magnetic field sources in a direction that differs from the azimuthal direction of the center of the range for the antenna which that static magnetic field source is associated with, by less than 20 degrees.
Optionally, the direction in which the expansion mechanism moves each of the at least two static magnetic field sources differs from the transverse component of the direction of magnetization (or the direction opposite to the direction of magnetization) of that static magnetic field source by a same angle, to within ±20 degrees.
In an embodiment of the invention, the probe is adapted to be inserted into a lumen of inner diameter greater than a minimum size, and when the imaging probe is inserted into a lumen of inner diameter twice the minimum size and the expansion mechanism is in its expanded state, the at least two static magnetic field sources and their associated antennas are close enough to the wall of the lumen so that at least part of the sub-region of each of their associated antennas is inside the wall.
Optionally, at least 40% of the NMR signal power received by said associated antennas originates from excited nuclei inside the wall.
Optionally, the parts of said sub-regions within the wall cover a set of azimuthal angles around the wall that does not have any gap greater than 90 degrees.
Optionally, the set of azimuthal angles around the wall does not have any gap greater than 45 degrees.
Optionally, the set of azimuthal angles around the wall covers more than 180 degrees.
Optionally, the set of azimuthal angles around the wall covers 360 degrees.
Optionally, the parts of said sub-regions within the wall cover said set of azimuthal angles within a longitudinal range of less than 15 mm.
In an embodiment of the invention, the time-varying magnetic field that the antenna associated with at least one of the static magnetic field sources creates is predominantly a dipole field outside the imaging probe.
Optionally, said antenna comprises two coils, adjacent to opposite sides of said static magnetic field source, which two coils run in phase with each other, and the time-varying magnetic field that said antenna creates in the center of the sub-region of said antenna is primarily transverse to the longitudinal axis.
Optionally, said antenna comprises a coil which wraps around said static magnetic field source longitudinally, and the time-varying magnetic field that said antenna creates in the center of the sub-region of said antenna is primarily transverse to the longitudinal axis.
Optionally, the dipole field has a dipole moment oriented at an angle greater than 45 degrees from the longitudinal axis.
Optionally, the static magnetic field that said static magnetic field source creates is predominantly a dipole field outside the imaging probe, and the dipole moment of the static magnetic field is oriented at an angle greater than 45 degrees from the dipole moment of the time-varying magnetic field.
Optionally, the dipole moment of the static magnetic field is oriented at an angle greater than 45 degrees to the longitudinal axis.
In an embodiment of the invention, the probe includes an expansion mechanism with a retracted state and an expanded state, which mechanism, when it expands, moves at least two of the static magnetic field sources in different directions transverse to the longitudinal axis.
In an embodiment of the invention, the probe is adapted to be inserted into a lumen of inner diameter greater than a minimum size, and, when the imaging probe is inserted into a lumen of inner diameter twice the minimum size and the expansion mechanism is in its expanded state, the probe presses against the wall of the lumen with sufficient force to stabilize the position of the probe sufficiently so that relative motion of the probe and the wall does not substantially affect the image quality.
Optionally, the probe is adapted to be inserted into an artery, and, when the lumen is an artery, the probe presses against the wall with no more than one atmosphere of pressure.
Optionally, the expansion mechanism comprises an expanding basket mechanism.
Alternatively or additionally, the expansion mechanism comprises an expanding helical mechanism.
Optionally, the expansion mechanism comprises a shape memory alloy.
Optionally, the expansion mechanism expands from the collapsed state to the expanded state when the temperature of the shape memory alloy is raised from below to above a shape memory transition temperature.
Alternatively or additionally, the shape memory alloy is in a superelastic state.
In an embodiment of the invention, the expansion mechanism comprises a distal end and a proximal end, and the expansion mechanism expands from the collapsed state to the expanded state when the distal end and the proximal end are brought closer together.
Alternatively or additionally, the expansion mechanism comprises a balloon, and the expansion mechanism expands from the collapsed state to the expanded state when the balloon is expanded.
Optionally, the plurality of static magnetic field sources comprises two static magnetic field sources.
Optionally, the plurality of static magnetic field sources comprises three static magnetic field sources.
Optionally, the plurality of static magnetic field sources comprises four static magnetic field sources.
Optionally, the plurality of static magnetic field sources comprises more than four static magnetic field sources.
In an embodiment of the invention, the sub-regions together have a longitudinal extent greater than 20% of the length of the probe in the longitudinal direction.
Optionally, the sub-regions together have a longitudinal extent greater than 50% of the length of the probe in the longitudinal direction.
Optionally, the sub-regions together have a longitudinal extent greater than 2 mm.
Optionally, the sub-regions together have a longitudinal extent greater than 5 mm.
Optionally, the sub-regions together have a longitudinal extent greater than 15 mm.
Optionally, the sub-regions together have a longitudinal extent greater than 30 mm.
In an embodiment of the invention, at least one of the static magnetic field sources is a permanent magnet element in the shape of a cylinder with a piece sliced off, the plane of the slice being within 20 degrees of parallel to the axis of the cylinder, the permanent magnet being magnetized in a direction substantially perpendicular to the axis of the cylinder and parallel to the plane of the slice.
There is further provided, in accordance with an exemplary embodiment of the invention, a probe with a longitudinal axis, for use in an NMR system, the probe comprising:
Optionally, the static magnetic field sources arranged in the row comprise three magnetic field sources.
There is further provided, in accordance with an exemplary embodiment of the invention, a probe with a longitudinal axis, for use in an NMR system, the probe comprising:
Optionally, the static magnetic field sources arranged in the row comprise four magnetic field sources.
Optionally, the at least one antenna comprises a plurality of coils, one for each static magnetic field source in the row, that is not located at an end of the row.
Optionally, each of the coils in the plurality of coils is located on a same side of the probe, adjacent to a different one of the static magnetic field sources in the row, that is not located at an end of the row.
Optionally, the at least one antenna comprises a coil.
Optionally, at least two of the static magnetic field sources in the row touch each other.
Alternatively or additionally, at least two of the adjacent static magnetic field sources in the row are separated by a gap.
Optionally, the gap at its narrowest point is smaller than 20% of the largest diameter of the probe at the gap.
Optionally, the plurality of static magnetic field sources comprise a plurality of permanent magnets.
There is further provided, in accordance with an exemplary embodiment of the invention, a probe for use in an NMR system, the probe comprising:
Optionally, the permanent magnet element is in the shape of a hollow right circular cylinder with the piece sliced off, and the slice extends into the hollow part of the cylinder, thereby making the permanent magnet element C-shaped.
Optionally, the probe is adapted to be inserted into a lumen, and includes a balloon which fits into the volume of the removed slice when the balloon is in a deflated state, and which holds the imaging probe against the wall of the lumen when the balloon is in an inflated state.
Optionally, the antenna is located on a different side of the cylinder than the slice.
Optionally, the cylinder is a right circular cylinder, the permanent magnet element has a slot carved into the side of the cylinder where the antenna is located, and the antenna is located in the slot, thereby confining the antenna substantially to the envelope of the cylinder.
Optionally, the permanent magnet element is in the shape of a cylinder with two pieces sliced off, the plane of each of the two slices being within 20 degrees of parallel to the axis of the cylinder.
Optionally, the two slices are within 20 degrees of parallel to each other. Optionally, the two slices are on different sides of the cylinder.
Optionally, the cylinder is a right circular cylinder.
Optionally, the probe includes an electrical component associated with the antenna, which component is located outside the surface of the slice, and within the cylinder.
In an embodiment of the invention, the plurality of static magnetic field sources comprise a permanent magnet, with substantially uniform cross-section transverse to the longitudinal axis, magnetized substantially uniformly in a direction substantially perpendicular to the longitudinal axis, and including at least one end cap, located at one end of the permanent magnet, sufficiently thick and permeable to make the magnetic field at a distance ⅔ of the magnet radius beyond the outer surface of the magnet vary by less than 10% longitudinally between the center of the magnet and a point ⅘ of the magnet radius away from said end of the magnet.
There is further provided, in accordance with an exemplary embodiment of the invention, a probe for use in an NMR system, the probe comprising:
Optionally, the at least one end cap is sufficiently thick and permeable to make the magnetic field at a distance ⅔ of the magnet radius beyond the outer surface of the magnet vary by less than 10% longitudinally between the center of the magnet and a point ⅘ of the magnet radius away from said end of the magnet.
Optionally, the at least one end cap comprises two such end caps, located at each end of the permanent magnet.
Optionally, the at least one end cap has a thickness at least equal to one tenth of the diameter of the permanent magnet in the direction of magnetization.
Optionally, the time-varying magnetic field differs in direction from the static magnetic field by more than 60 degrees and less than 120 degrees, somewhere in the sub-region.
Optionally, at least one of the static magnetic field sources comprises a material with skin depth greater than the largest dimension of said static magnetic field source, at the proton nuclear resonance frequency at the maximum static magnet field in the region outside the probe.
Optionally, at least one of the static magnetic field sources comprises sintered material.
Optionally, at least one of the static magnetic field sources comprises ferrite.
Optionally, the probe is an imaging probe, and the NMR system is an MRI system.
Optionally, the one or more antennas comprise a single antenna capable of creating the time-varying magnetic field, and receiving the NMR signals and generating the NMR electrical signals.
Alternatively or additionally, the one or more antennas comprise:
There is further provided, in accordance with an exemplary embodiment of the invention, an NMR system comprising a probe according to an embodiment of the invention, a power supply which transmits power to at least one of the antennas of the probe to create the time-varying magnetic field, and a data analyzer which reconstructs NMR characteristics of material in the sub-region from the NMR electrical signals generated by at least one of the antennas of the imaging probe.
Optionally, all of the at least one antennas that the power supply transmits power to are different from all of the at least one antennas that generate the NMR electrical signals from which the data analyzer reconstructs the NMR characteristics.
Alternatively, at least one of the at least one antennas both creates the time-varying magnetic field and generates the NMR electrical signals from which the data analyzer reconstructs the NMR characteristics.
Optionally, the NMR system is an MRI system, the probe is an imaging probe, and the data analyzer comprises an image reconstructor which reconstructs an image.
There is further provided, in accordance with an exemplary embodiment of the invention, an NMR system comprising:
There is further provided, in accordance with an exemplary embodiment of the invention, a non-imaging NMR system, comprising:
Exemplary embodiments of the invention are described in the following sections with reference to the drawings. The drawings are generally not to scale and the same or similar reference numbers are used for the same or related features on different drawings.
MRI Probe with Obliquely Magnetized Magnets
This probe is similar to that described by Blank et al, in U.S. Pat. No. 6,704,594, except for the direction of magnetization of the magnets, indicated in
Probe 100 has a field of view 112, bounded approximately in x and z by the box shown in
In
The magnetic field around probe 100 in
Optionally, the direction of magnetization of the two magnets is oriented at an angle different from 45 degrees to the longitudinal axis. The optimal angle to maximize the magnetic field about the field of view of the probe, depends on the relative length, diameter and gap distance of the magnets, as well as on where the desired volume of view is located. The optimum angle is optionally determined by calculating the magnetic field about the field of view, for different angles of magnetization of the magnets, using magnetic finite element software. In some cases, the optimal angle may depend not only on maximizing the magnetic field about the field of view, but also on constraints on the magnetic field gradient, which also depends on the angle of magnetization of the magnets. For example, if the RF circuitry is only capable of handling a certain maximum bandwidth, then it may be desired to have a more uniform magnetic field in the field of view, even at the cost of having a lower magnetic field. Or, it may be desirable to have a greater magnetic field at a certain distance from the magnets, to allow the field of view to extend out that far, even at the cost of having less than the greatest possible magnetic field close to the magnets.
Optionally, there is no gap 116 between magnets 102 and 104, but the magnets are bonded together. However, there is a potential advantage to having a gap between the magnets, particularly right behind RF coil 108, where the RF field is greatest. A strong RF field can produce significant eddy currents in the magnets, as well generating sound waves in the magnets at the RF frequency, a phenomenon called magnetoacoustic ringing. Eddy currents and sound waves can both cause heating of the magnets, possibly raising their temperature too high, and wasting RF power, reducing the RF power available for generating MRI signals. Even if the magnets are almost touching each other, having at least a small gap between them, so the magnets are not in electrical contact with each other, or not in mechanical contact with each other, may reduce eddy currents and magnetoacoustic ringing.
Methods of Operation of MRI Probes
The method of operation of probe 100 is briefly described here with reference to
When told to do so by controller 122, power supply 124 transmits RF power through transmission channel 132 to interface 128. Optionally, the RF power is in the form of pulse sequences known in the art, for example a CPMG spin echo sequence. A catheter 134 carries transmitted RF power from interface 128 to probe 100, which was previously placed in blood vessel 136 using techniques known in the art, or described herein or in the related application filed concurrently, and the RF power excites nuclei in field of view 112, shown in the wall of blood vessel 136. Alternatively, probe 100 is placed in other body cavities, or used outside the body. The excited nuclei emit NMR signals, which are received by probe 100 and carried by catheter 134 from probe 100 to interface 128.
Preferably, for safety reasons, there is an over-power protection circuit, located for example in controller 122 or interface 128, which limits the RF power, averaged over an appropriate time interval, which can be delivered to probe 100. If the average RF power is too high, then the power supply is turned off, and/or the power supply is disconnected from the catheter, and/or other measures are taken to stop the further transmission of RF power to the probe, and optionally an alarm sounds.
Optionally, for safety reasons, interface 128 uses optical coupling to transmit control signals from controller 122 to catheter 134, or to any part of interface 128 that is in electrical contact with catheter 134, and to transmit data signals from catheter 134 to controller 122.
Probe 100 includes magnets and an RF coil for transmitting and receiving, as shown in
The NMR signals received by interface 128 are conveyed, optionally after being amplified and/or otherwise processed by interface 128, through receiving channel 138 to controller 122, where the data is optionally used to reconstruct an image of the blood vessel wall, showing, for example, the presence or absence of plaque. Additionally or alternatively, the data is used for other purposes, for example for NMR spectroscopy, or to determine an average T1, T2, and/or diffusion rate for the blood vessel wall in the vicinity of the probe. Optionally, the reconstructed image is displayed in real time on a monitor 140. Optionally, the displayed image, or spatially averaged non-imaging information, combines data received at different times from different fields of view as the probe is turned or moved. The MRI system and method outlined in
Probe 100 optionally uses one or more of several methods to obtain resolution in the radial, azimuthal and longitudinal directions, in producing an image. The static magnetic field falls off with increasing radial distance from the probe and also with increasing azimuthal angle away from the x-axis, as shown by contours 118 in
Optionally, probe 100, or any of the other MRI probes shown in the drawings, uses a high RF bandwidth, and a large number (up to thousands) of spin echoes, in order to obtain a reasonably high SNR with a high field gradient. Alternatively, the probes have an external region of low field gradient, and a more conventional MRI pulse sequence is used, with smaller RF bandwidth and fewer echoes. A potential advantage of using high field gradient in the imaging region, as discussed above, is that the field can be higher in the imaging region, and the imaging region can be broader.
For example, for intravascular MRI probes designed to be used in blood vessels as small as 2 mm in diameter, the field gradient is optionally as great as 50 tesla/meter, or 100 tesla/meter, or 150 tesla/meter, or 200 tesla/meter, or 300 tesla/meter, or 400 tesla/meter, or even higher. The bandwidth is optionally as great as 0.25 MHz, or 0.5 MHz, or 0.75 MHz, or 1 MHz, or 1.25 MHz, or 1.5 MHz, or 2 MHz, or even greater. Note that 1 MHz bandwidth corresponds to 0.024 tesla, so, for example, if the bandwidth is 1 MHz and the gradient is 200 tesla/meter, then the slice has a radial thickness of 0.12 mm. With different values of field gradient and bandwidth, the slice radial thickness is, for example, 0.03 mm, or 0.06 mm, or 0.2 mm, or 0.3 mm, or greater or less than these values.
Another potential advantage of using a large field gradient is that it is easier to measure the diffusion rate of molecules in the tissue being imaged, which can be used to distinguish healthy tissue from the necrotic lipid-rich core in vulnerable plaque. A higher diffusion rate, in the presence of sufficiently strong field gradient, causes the tissue to appear as though it has a shorter effective T2 value, which depends on the diffusion rate. This effective T2 value, called Tc, may be more useful that the actual T2 value for distinguishing vulnerable plaque from healthy vascular tissue. For example, T2 is typically 30 to 50 milliseconds in vascular tissue, and diffusion rates range from 1.6×10−9 m2/sec in healthy tissue to 0.4×10−9 m2/sec in the necrotic core. To have Tc depend primarily on the diffusion coefficient rather than on T2, for this range of diffusion coefficients and T2, and for a bandwidth of about 1 MHz, the field gradient should be greater than about 100 T/m. Although diffusion of this magnitude can also be measured with smaller field gradients and smaller bandwidth, using fewer echoes of longer duration, the duration and time between echoes should preferably not be much greater than T2, and in conventional MRI there is a limit to how strong a field gradient can be used, since the gradient is supplied by gradient coils, rather than by the main magnet. Consequently, it is difficult to measure diffusion rates much smaller than 10−9 m2/sec in conventional MRI. With a self-contained intravascular MRI probe with high static magnetic field gradient and using high bandwidth RF pulses, it is quite possible to measure diffusion rates that are smaller than this by an order of magnitude or more.
MRI Probe with Multiple Sub-Probes
For each sensor, the RF coil is located against one side of the magnet, centered halfway between the north pole side and the south pole side of the magnet, so the azimuthal locations of the coils also differ by 90 degrees between one sensor and the next one below it. Alternatively the coils are located at a different location against the side of each magnet. However, centering the RF coil halfway between the north pole and south pole of each magnet maximizes the NMR signal, which depends on the component of the RF magnetic field that is perpendicular to the static magnetic field.
An alternative to the RF coil configuration in
Alternatively, one or more of the sensors in
Preferably, basket structure 304, or whatever expansion mechanism is used, does not press so hard against the blood vessel wall that it might damage plaque, which can be dangerous. For example, the expansion mechanism does not exert a pressure of more than one atmosphere on the blood vessel wall. The field of view for each sensor in probe 200 has a limited azimuthal extent, centered around the RF coil. For example, for each sensor, the field of view is 75 degrees, measured from the center of that sensor. If, as in
Optionally, basket structure 304 is made of a shape memory alloy such as NiTi, and it is expanded by heating it above its transition temperature. Alternatively or additionally, it is made of superelastic shape memory alloy, or it is made of a material other than shape memory alloy, and mechanical means are used to expand the basket structure. For example the two ends of the structure are pulled toward each other by a wire that runs through a catheter, causing the basket structure to bow outward. Alternatively or additionally, a balloon is used to expand the basket structure, although in that case the basket structure may not have the potential advantage of not occluding blood flow.
An expanding structure suitable to use in place of basket structure 304 is described in more detail in the related patent application filed concurrently.
Although the field of view for each sensor in
MRI Probes with Longitudinally Magnetized Magnets
The probe shown in
The larger number of magnets in
Optionally, in probe 600 or in a similar probe with more magnets and RF coils, images or other data obtained from the NMR signals received by the different coils are simply averaged together. Alternatively, the NMR signals received by different coils, or by different sets of coils, are not combined together, but are recorded separately and analyzed to obtain finer longitudinal resolution of the field of view.
MRI Probes With Other Magnet Configurations
Alternatively, instead of the longitudinal portions of the RF coils being connected by conductors that go around the circumference of magnet 702 azimuthally, as shown in
MRI Probes with End Caps on Magnets
A probe such as probe 1200 or probe 1300, with magnet magnetized perpendicular to the longitudinal axis, would, if it were infinitely long, produce a magnetic field that is perpendicular to the longitudinal axis everywhere, and independent of axial position. But, with a probe of finite length, the flux lines near the end of the magnet will tend to bow out axially, reducing the field strength, at a given radius and azimuth, near the ends of the magnet. This effect may be significant within one or two magnet diameters away longitudinally from the ends.
With end caps of sufficient thickness and sufficiently high saturation flux density, much of the flux originating near the ends of the magnet will go through the end caps, instead of bowing out axially past the ends of the magnet. This results in a magnetic field that is substantially more uniform, as a function of axial position, for a given radius and azimuth over most of the length of the magnet, and which falls off substantially more abruptly near the ends of the magnet. For example, the magnetic field at a distance ⅔ of the magnet radius beyond the outer surface of the magnet varies by less than 10% longitudinally between the center of the magnet and a point ⅘ of the magnet radius from the end of the magnet.
This may be seen in
The minimum thickness at which the end caps are effective depends on the saturation flux density of the end cap material (typically about 2 tesla for iron or steel), and on the remanence of the permanent magnet material (typically about 1.4 tesla for high quality rare earth magnets), and on the diameter of the magnet. For example, the end caps have a thickness at least 5% of the magnet diameter, or at least 10% of the magnet diameter, or at least 20%, or at least 40%. The probe shown in
Optionally, instead of using an end cap to achieve a more uniform magnetic field and a more abrupt fall-off at one or both ends of the magnet, a similar result is achieved by making the magnet thicker toward one or both ends. Magnetic finite element software may be used to calculate the shape of the magnet that would produce the desired magnetic field.
The more uniform magnetic field and abrupt fall off at the ends may have potential advantages. Radial resolution in these probes may be achieved by exciting nuclei in the imaging region at different RF frequencies, corresponding to the magnetic resonance frequency at different field strengths. If the contours of constant field strength are at nearly constant radius, as in
Manufacturing Process
Regardless of whether the probe comprises multiple sub-probes, and regardless of whether there is more than one magnet or more than one coil in the probe or in each sub-probe, and the direction of magnetization of the magnet or magnets, the same basic procedure is optionally used in assembling the probe or sub-probes. An exemplary procedure may be summarized as follows:
1) Assemble “short” probe (or sub-probe).
2) Assemble electric circuit.
3) Assemble full probe (or sub-probe).
4) Assemble proximal part of catheter.
5) Join catheter to probe (or to each sub-probe).
6) Assemble optional expansion mechanism (e.g. balloon, or basket) on probe or sub-probes.
The “short” probe (or sub-probe), in the terminology used by Topspin Medical, Inc., comprises a magnet or magnets, a structure for mounting the probe on a guide wire or basket, one or more coils, and one or more capacitors (often varicap diodes) for RF tuning, typically in series with a coil. The full probe (or sub-probe) comprises a short probe and an electric circuit, connected with coaxial cables, and a structure joining them together.
Once tube 904 is bonded to groove 902, two coaxial cables 906 and 908 are laid through groove 902 on top of tube 904, and bonded in place against tube 904, for example with cyanoacrylate. For a design of probe 100 with a diameter of 5.5 French (1.83 mm), magnets 102 and 104 are 1.6 mm in diameter, and the coaxial cables are 70 micrometers in diameter. Optionally, the coaxial cables have spiral wrap shielding rather than braided shielding, and leads of cables 906 and 908 are easily stripped by untwisting the shielding. The shielding of the two cables is optionally bonded together with conductive epoxy.
The lead of cable 906 is then soldered to a first pad of a variable capacitor 910, for example a varicap diode whose capacitance can be adjusted by applying a DC voltage to it, for RF tuning, and the soldered bond is optionally strengthened by applying cyanoacrylate. Optionally, cyanoacrylate is used immediately after soldering to strengthen one or more of the soldered bonds in the probe. Variable capacitor 910 is then placed in a gap 912 between the two magnets, and UV curable glue is optionally placed between the variable capacitor and the magnets. Alternatively, another kind of glue is used. Using a 40× microscope, the variable capacitor is optionally then maneuvered precisely into position, for example within 10 micrometers of the magnet surface, and a UV light source is used to cure the UV glue. In the case of probe or sub-probe designs with a single magnet, or with no gap between magnets, variable capacitor 910 is optionally placed in a slot 106 on the front of the magnets (i.e. on the same side of the probe as the field of view), where an RF coil 108 is also located. For example, variable capacitor 910 is placed behind coil 108, or to the side of coil 108. Coil 108 is then placed in slot 106, and optionally is precisely positioned using UV curable glue, or another kind of glue, in the same way as variable capacitor 910 is positioned. Having variable capacitor 910 close to the coil has the potential advantage of reducing stray capacitance and stray inductance in the leads connecting them, improving the efficiency of the RF coupling to the field of view of the probe, and assuring that the probe is tuned to an intended frequency range. Once the variable capacitor and coil are in place, the stripped lead of coaxial cable 908 is soldered to one lead of coil 108, and the second pad of variable capacitor 910 is soldered to the other lead of coil 108. Gap 912 between the magnets, and slot 106, are then optionally filled in with epoxy, forming a smooth cylindrical surface continuous with the cylindrical surface of the magnets.
Optionally, the short probe is then covered with a thin film of vapor deposited aluminum, except for the coaxial cables which are masked, since they are already shielded, in order to shield the RF coil from far-field noise, without shielding out the near-field signal from the imaging region, and without interfering with the transmission of RF power to the near-field imaging region. Heat shrink, for example made of PET, is then optionally placed around the probe and shrunk, further protecting it and preventing any relative movement between the different components of the probe.
To manufacture magnets 102 and 104, for the 5.5 French diameter probe shown in
To manufacture coil 108, the coil is optionally first wound between two flat chrome-coated plates, one of which has a hard steel core, optionally using 20 micrometer insulated copper wire with bonding material on the outside. The coil is then optionally heat-treated to melt the bonding material and to bond the turns together. The coil is then optionally dipped in acrylic and baked, and optionally bent to conform to the shape of the cylinder, as shown in
Optionally the coil is wound with wire that is thicker or thinner than 20 micrometers in diameter. However, using wire about 20 micrometers or less in diameter has the potential advantage that the wire diameter is less than the skin depth in copper at the RF frequencies of interest, so that the current penetrates well into all of the wire cross-section. Using wire about 20 micrometers or more in diameter has the potential advantage that it is easier to wind the coil than if thinner wire were used. And using wire about 20 micrometers in diameter has the potential advantage that, for a coil of these dimensions, the coil resistance and inductance have convenient values for impedance matching to off-the-shelf RF amplifiers and power supplies. For example, with 20 micrometer diameter wire, the coil DC resistance is about 25 ohms and the inductance is about 18 microhenries. Alternatively, using different diameter wire or different coil dimensions or a different fill factor or a different material for the coil, the resistance is, for example, about 5 ohms, or 10 ohms, or 50 ohms, or 100 ohms, or greater or less than these values. Alternatively, the inductance is about 3 microhenries, or 8 microhenries, or 40 microhenries, or 100 microhenries, or greater or less than these values. The RF field per ampere produced by the RF coil in the imaging region ranges from about 0.05 tesla/amp to 0.15 tesla/amp. Alternatively, the RF field per amp is about 0.01 tesla/amp or 0.03 tesla/amp or 0.30 tesla/amp, or greater or less than these values.
To assemble electric circuit 1006, as shown in
Once the short probe and the electric circuit have been assembled, the full probe, comprising the short probe, the electric circuit, coaxial cables 906 and 908 connecting them, guide wire tube 904, and optionally a tube for inflating a balloon attached to the probe, are optionally encapsulated in a tube, for example a polyimide tube, which is filled with epoxy. Optionally, flex molding is optionally bonded to the proximal end of the short probe, and the shaft of the catheter is pulled over the electric circuit, and bonded to the flex molding. Alternatively, the shaft of the catheter is bonded directly to the proximal end of the short probe, but the flex molding may provide greater flexibility.
Optionally, in addition to electric circuit 1006 and the electric elements located in the short probe, there are other electric elements which may be located, for example, in a package at the proximal end of the catheter, outside the body. These other elements may include, for example, one or more amplifiers for amplifying the received RF signal. If the same RF coil is used for receiving and transmitting RF power, then optionally there are elements which isolate the receiving amplifiers from the RF coil when it is in transmitting mode, in order to avoid damaging or saturating the amplifiers. An example of a “duplexer” circuit which incorporates such elements is shown in
Optionally, the proximal end of the catheter has three branches, for the guide wile, the coaxial cables and ground wire, and the balloon inflating tube, if there is a balloon attached to the probe. The three branches are optionally sealed with epoxy. Optionally, the branch with the guide wire has a special pressure seal, with a side branch for injecting saline solution and heparin, and a swiveler to allow the probe to be rotated. Optionally, the coaxial cables and ground wire are wrapped a few times around the guide wire near the point where the three branches come together, in order to avoid straining the coaxial cables and ground wire when the catheter rotates.
RF Antenna with Short Dead Time
During the transmission, high voltage RF pulses coming from transmission line 1604 are coupled to RF coil 108 through a multi-inputs RF transformer 1606. Cables 1010 and 1012, running through the catheter, connect transformer 1606 to RF coil 108, and to the rest of a catheter circuit 1601, shown in more detail in
Optionally, the resistance of the primary turns of each of the toroid protectors is no more than a few ohms when the secondary circuit is open. Optionally, the apparent resistance of the primary turns is several hundred ohms when the secondary circuit is shorted out. With these values, circuit 1600 will not ring too long (extending the dead time of the LNA) after the RF transmitter is turned off, but the circuit will not dissipate too much power. Optionally, the inductance of the toroid protectors is not more than a few microhenries, which enables good impedance matching at the RF frequencies of interest, on the order of 10 MHz. Optionally, the junction voltage of the Schottky diodes is less than about 0.3 volts, and their capacitance is less than about 1 picofarad, to avoid excessive losses and to provide good coupling to the LNA.
The circuit shown in
The dead time of LNA 1602 is limited by the ringing of circuit 1600, including RF coil 108 and cables 1010 and 1012. A dead time of only 4.5 microseconds has been achieved using the circuit shown in
MRI Probe with Non-axisymmetric Balloon
Optionally, probe 100, or any of the other MRI probes shown in the drawings, has a non-axisymmetric elastically expandable balloon mounted on it. The balloon, not shown in the drawings, expands only or mostly on one side of the probe, where the balloon is more elastic, pressing the other side of the probe (where the balloon is more rigid) against the wall of the blood vessel, so that field of view 112 falls inside the wall. The balloon is made, for example, by placing a tube of elastic material around the probe, then masking one side of the tube, and depositing a stiffening material, such as parylene, to the outside of the tube optionally on all sides. When the masking is removed, one side of the tube will be coated with the stiffening material, and will be relatively rigid, while the other side will be free of the stiffening material in the region that was covered by the mask, and will remain elastic. Alternatively, the stiffening material is not deposited on all sides, but only on the side where it is needed, and in this case masking is optionally not used. Optionally, the uncoated part of the balloon is sufficiently thin and elastic so that it can be used in blood vessels with a large range of different diameters, in contrast to the relatively inelastic balloons typically used for expanding stents in arteries.
Thin Film RF Shielding
Optionally, probe 100, or any of the other MRI probes shown in the drawings, is coated with an RF shielding material. Optionally, the RF coil (for example RF coil 108 in probe 100) is coated with the shielding material. The shielding material is much thinner than a skin depth at the RF frequency, but has a conductivity very much greater than the RF frequency times the electrical permittivity of body tissues. The shielding material hardly affects the near-field RF fields that the probe uses to produce MRI images, but largely shields out interference from plane waves at the RF frequency produced by distant sources, much more than a wavelength away. The shielding comprises, for example, a layer of aluminum 300 nanometers thick, applied to the probe by vapor deposition. Optionally, in order to improve the adhesion of the aluminum to the probe or to the RF coil, an even thinner layer of titanium, for example 30 nanometers thick, is applied first.
General
The invention has been described in the context of the best mode for carrying it out. It should be understood that not all features shown in the drawings or described in the associated text may be present in an actual device, in accordance with some embodiments of the invention. Furthermore, variations on the method and apparatus shown are included within the scope of the invention, which is limited only by the claims. Also, features of one embodiment may be provided in conjunction with features of a different embodiment of the invention. As used herein, the terms “have”, “include” and “comprise” or their conjugates mean “including but not limited to.”
