KR20110104807A - Rf-resonator - Google Patents

Abstract

According to the present invention, the conventional clinical RF-Coil is suited to humans, so that MR signals are lost due to the filling factor when obtaining MR images from small objects such as small animals (Rat, mouse), thereby obtaining sufficiently good images. In order to reduce such signal loss, RF-Coil for receiving MR signal is also sufficiently miniaturized to increase filling-factor and to reduce overall resistance (R) by miniaturizing size to improve SNR in clinical MRI equipment. Small FOV (High Field of View) and High Resolution Image which could not be obtained in the past, small animal research using various animal models and image acquisition by organic materials are available. An animal high frequency resonator is provided for use in various fields of small animal MRI research.

Description

High Frequency Resonator for Animals {RF-Resonator}

FIELD OF THE INVENTION The present invention relates to nuclear magnetic resonance (hereinafter referred to as NMR) devices, and more particularly to high frequency (hereinafter referred to as RF) animal resonators useful for devices for transmitting or receiving RF signals.

Conventionally, NMR phenomena have been useful by chemists who study the molecular structure of micromolecules of organisms in vitro. Typically, NMR spectrometers useful for this purpose are designed to accommodate relatively small samples of the material to be studied. Recently, however, NMR has been developed as an image forming method useful for obtaining images of anatomical features of the human body, for example. Such an image depicting parameters associated with nuclear spin (typically proton hydrogen associated with tissues of water) may be a foreign diagnostic value that determines the health of the tissue in the area tested. NMR technology has also been extended to active spectroscopy of elements such as phosphorus and carbon, providing researchers with tools to study chemical processes in living tissues, for example. The use of NMR to produce images and research by spectroscopy of the human body requires the use of specially designed system components such as magnets, gradients and RF resonators.

As a background, the NMR phenomenon occurs in an atomic nucleus with an odd number of protons or neutrons. Due to the spins of protons and neutrons, a sample composed of such nuclei is placed in a homogeneous magnetic field B0 such that a number of nuclear magnetic moments are in line with the field to generate macroscopic net magnetization M in the direction of the field. Each such nucleus represents a magnetic moment. Under the influence of the magnetic field B0, the magnetic moment process, which is straight with respect to the axis of the field at frequency, depends on the strength of the applied magnetic field and the nature of the nucleus. Each precession frequency ω is also referred to as the rammer frequency given by the lamer equation ω = γB, where γ is the magnetic rotation rate (which is a constant for each NMR isotope), and B is the magnetic field performed by the nuclear spin (other field plus B0). It will therefore be clear that the resonant frequency depends on the strength of the magnetic field at which the sample is located.

The direction of the magnetization M normally induced along the magnetic field B0 may be disturbed by oscillating the magnetic field at or near the rammer frequency. Typically, such magnetic field designated B1 is applied at right angles to the direction of the B0 field by an RF pulse through a coil connected to a radio frequency transmitter. Under the influence of RF excitation, the magnetization M rotates in the direction of the B1 field. In NMR studies, it is typically desirable to apply RF pulses of sufficient magnitude and duration to rotate magnetization M into a plane perpendicular to the direction of the B0 field. The plane is generally referred to as the cross section. Under the stop of RF excitation, the nuclear moment is rotated into a precession of the cross section around the direction of the orbiting field. The vector sum of the spins is in the form of precessed bulk magnetization that can be detected by the RF coil. The signal sensed by the RF coil, or NMR signal, is characterized by the particular chemical environment and magnetic field in which the nucleus is located. In NMR imaging applications, the NMR signal is observed in the presence of magnetic field gradients that make encode spatial information useful in the signal.

This information is later used to some extent to reconstruct the image of the researched purpose well known to those familiar with the technique.

Many NMR systems approach the sample region along the Z-axis of the Cartesian coordinate system along the direction of the orbiting magnetic field that will typically be selected. This relates to an NMR system using a superconducting magnet, which is considered for efficiency because NMR imaging systems are used throughout all humans. Therefore, the sample or human body is inserted along the direction of the static field, and the RF transmit and receive coils are often wound on a cylindrical coil that forms an axis parallel to the static field. The RF magnetic field should be perpendicular to the idle RF magnetic field and therefore perpendicular to the cylindrical coil.

1A and 1B show a saddle type RF coil of conventional design. The coil consists of a single rotation 1 and 3 connected in parallel and is driven at points 7 and 9 at both ends of the tuning capacitor 8. Such coils are typically formed from copper tubing 5 installed on a non-conductive (high insulator) cylinder 11 as shown in FIG. 1B. Each coil rotation is dimensioned over 120 mm 3 of the circumference of the cylinder.

The coil areas of the connections 7 and 9 are dimensioned over approximately 60 mm of circumference. For maximum RF field uniformity, the side of the coil parallel to the longitudinal axis of the cylinder must be equal to the two cylinder diameters (D). However, coils having two diameters of lateral length are impractical because RF energy is located at the site of the patient. Therefore, in fact the side length of the coil is reduced to approximately one diameter length.

FIG. 1C is another embodiment of a conventional RF coil similar to that shown in FIG. 1A but driven at points 19 and 20 across the coil rotations 15 and 17 and capacitor 18 connected in series. It is shown. The coil shown in FIG. 1C is typically used for NMR studies of the head.

In the case of each of the coils shown in Figs. 1A and 1C, the RF field generated is somewhat non-uniform and therefore undesirable for NMR image forming applications.

Therefore, it is clear that the current distribution needs to be controlled by multiple coil windings to produce a uniform B1 field. In addition, as already described previously, the coil geometry is naturally free to approach along its longitudinal axis for positioning of the patient or other NMR sample. The B1 field must also be perpendicular to the cylindrical axis of balance to be selected parallel to the direction of the B0 field.

An improved RF coil, claimed in the related patent application and described in more detail, is shown in FIGS. 2A and 2B. Referring first to FIG. 2A, it is an eight-section typical embodiment consisting of conductive loop elements 23 and 25 separated into two spaces interconnected by eight axes of conductive elements 27 to 34. An RF coil is shown. Each conductive element provides at least one capacitive element, such as the indicated 35-42, that matches the capacitor in each of the elements 27-34. Conductive loop elements 23 and 25 each have eight series connected conductive elements in the loop cutoff (in an 8-section coil) between conductors of adjacent axes. The conductive element, indicated generally by the reference numeral 45 in the loop 25 and 43 in the loop 23, represents the distributed intrinsic inductance in the conductor consisting of the loop element. The inductance is necessary for proper coil operation to achieve the desired phase shift. The RF coil is excited by applying an RF energy source (not shown) to leads 20 and 22 across the input capacitor at one of the conductive elements (such as capacitor 35).

In the overall implementation of the NMR study, information can be found to be advantageous in increasing the intensity of the homogeneous magnetic field B0. This is desirable in the case of quantum imaging to increase the signal-to-noise ratio of the NMR signal. However, in the case of spectroscopy, this is necessary because any chemical elements studied (eg, phosphorus and carbon) are relatively deficient in the body so as to be needed to detect useful signals of high magnetic fields. As is evident from the rammer frequency, the increase in magnetic field B entails a corresponding increase in frequency, thus causing an increase in the predetermined resonance frequency of the transmitter and receiver coils. This complicates the design of an RF coil large enough to accommodate the human body. It is one of the difficulties that the RF field generated by the coil must be equal for the part of the body to be studied in order to generate more consistent measurements and images. The generation of a uniform RF magnetic field over a large volume is due to the consequences of unwanted resonant capacitance between different parts of the RF coil and between the RF coil and the NMR sample itself, which limits the highest frequency at which the surrounding object or coil can create resonance. It becomes even more difficult at high frequencies.

Conventional RF coils use coils of one turn or two turns in parallel to minimize inductance and increase the resonance frequency. An example of a conventional two-turn coil is the so-called " saddle coil " which is described in detail later. The concentration of resonance currents at such few rotations reduces the homogeneity of the B1 field and the sensitivity homogeneity for signals generated at different parts of the sample region. In addition, a lack of balance between the position of the tuning capacitor and the idle capacitance of the single rotating coil causes a non-uniform current distribution in the coil and a corresponding reduction in uniformity of field and signal sensitivity.

The present invention has been made to solve the above problems, the object of the present invention, therefore, the object of the present invention can generate a substantially homogeneous B1 field and RF having a practical uniform signal sensitivity over the area of interest To provide a coil.

It is a further object of the present invention to provide an NMR RF coil which exhibits an enhanced signal-to-noise ratio and which is feasible to generate a circularly polarized field.

It is still another object of the present invention to provide an NMR RF coil having tuning capacitances and currents distributed at many rotations.

According to the present invention, there is provided an NMR radio frequency coil having a pair of conductive loop elements that maintain a constant distance along a common longitudinal axis. Each loop element includes a plurality of continuously connected capacitive elements that are spaced at regular intervals along the circumference of the loop. Multiple layered conductive elements connect electrically conductive loop elements to each other in terms of adjacent elements of an electrically, continuously connected capacitive element. In a "high pass" embodiment of the RF coil, the conductive portion of the layering may be a wire, conducting tube or flat conductive tape with the inductance inherent required for proper coil operation.

The "band pass" embodiment of the coil is recognized by including a capacitive element in each conductive portion on the axis.

Features of the invention that are considered new are described in detail in the appended claims. Hereinafter, the configuration, method of operation, object, and advantages of the present invention will be described in detail with reference to the accompanying drawings.

To achieve these and other advantages and in accordance with the purpose of the present invention,

The RF coil shown in FIG. 2A is based on the lumped element slow wave delay line structure shown in FIG. 2B, where the same parts are denoted by the same reference numerals.

Referring now to FIG. 2B, the conductive loop elements 23 and 25 are represented by a continuous conductor having a set of conductive elements 43 and 45 connected in series. Similar inductance, of course, is not shown in FIG. 2B, but is also related to the conductive portions of elements 27 to 34. In general, the inductance associated with conductive elements 27-34 exhibits a smaller effect than the capacitive effect associated with separate capacitive elements 35-42 associated with each element. The RF coil shown in FIG. 2A is formed by electrically coupling the ends P-P 'and Q-Q' of the respective conductors 23 and 25 (FIG. 2B).

The RF coil shown in FIG. 2A and the corresponding lumped element network shown in FIG. 2B will pass a low frequency signal along the conductive element while a high frequency signal will be shorted by the capacitor and interrupted by the inductor. It is referred to as "low pass" because it tends to. Coils with multiple axes of conductors above and below 8 are also possible and include coils with odd or even axes of conductors.

According to the present invention, the conventional clinical RF-Coil is tailored to humans, so when MR images are obtained from small targets such as small animals (Rat, mouse), the MR signals are lost due to the filling factor so that a good enough image can be obtained. Could not be obtained. In order to reduce such signal loss, RF-Coil for receiving MR signal is also sufficiently small to increase filling-factor and size to reduce overall resistance (R) to improve SNR. Field of view (FOV) and High Resolution Image were obtained. Thus, it is possible to study small animals using various animal models and to acquire images according to organic, which can be used for small animal MRI research in various fields such as research on targeting specific areas using substances such as contrast medium.

1a is a diagram showing a conventional two-turn NMR RF coil connected in parallel used for the overall study.
FIG. 1B is a view in which the coil shown in FIG. 1A is provided in a cylindrical shape. FIG.
1C illustrates another form of a conventional series connected two-turn NMR RF coil, for example, used for NMR testing of the head.
2A is a perspective view of an eight-section low pass RF coil;
FIG. 2B is a perspective view showing a concentrator, low pass, low speed and delay line structure used in the coil shown in FIG. 2A; FIG.
FIG. 2C is a graph showing the frequency responsive to circumferential inductors in a circulated current versus low pass RF coil. FIG.
2d is a perspective view of an eight-section high pass RF coil.

Looking at the configuration of the high frequency resonator device of the present invention in detail as follows. In describing the present invention, detailed descriptions of related well-known functions or configurations are omitted in order not to obscure the subject matter of the present invention.

Referring now to FIG. 2B, the conductive loop elements 23 and 25 are represented by a continuous conductor having a set of conductive elements 43 and 45 connected in series. Similar inductance, of course, is not shown in FIG. 2B, but is also related to the conductive portions of elements 27 to 34. In general, the inductance associated with conductive elements 27-34 exhibits a smaller effect than the capacitive effect associated with separate capacitive elements 35-42 associated with each element. The RF coil shown in FIG. 2A is formed by electrically coupling the ends P-P 'and Q-Q' of the respective conductors 23 and 25 (FIG. 2B).

The RF coil shown in FIG. 2A and the corresponding lumped element network shown in FIG. 2B will pass a low frequency signal along the conductive element while a high frequency signal will be shorted by the capacitor and interrupted by the inductor. It is referred to as "low pass" because it tends to. Coils with multiple axes of conductors above and below 8 are also possible and include coils with odd or even axes of conductors.

FIG. 2C shows the frequency (shown along the horizontal axis in NHz) versus the conductive loop (23 in FIG. 2A or 25) to be obtained in the low pass RF coil of the 16 cut embodiment similar to the 8 cut coil shown in FIG. 2A. Shows a predicted graph of the current (shown along the vertical axis in AMPs) through one of the conductive elements). In order to predict the current, it is natural that the coil is driven with an RF source across the inductor in either the loop element 23 or 25 in FIG. 2A. Each of the eight current peaks 52 to 59 in FIG. 2C represents two resonance modes of the RF coil. Current peak 52 corresponds to two quadrature resonance modes that occur at approximately 64 MHz (protons at the Rammer frequency in 1.5 Tesla B0 fields) of particular interest for NMR imaging applications. The excitation of the coil at this frequency using the appropriate source of RF energy (not shown)

Will occur in proportional to the loop elements 23 and 25, where K = 1 to 16 represent the conductive element number starting at the input inductor. The current through the conductive element of the axis,

Will be changed sinusoidally, where K = 1 to 16 are as described above. The second resonance mode can be excited by activating the RF coil at the second input inductor at the right angle at the first.

The RF coil of the present invention can be assembled using any hypothesis technique. Head and body coils, each with 16 axes of conductive elements (16 cuts), are hypothesized for use at about 127.68 MHz. The head coil is hypothesized on a plastic or glass coil shape having an outer diameter of 0.28 m and a length of 0.4 m. In the head coil, the axially conductive element is also a copper alloy printed circuit board, or may be made from wires, but rather is formed from a thin, conductive copper tube. Suitable copper tubing is made to have an outer diameter of 3 mm (1/8 inch). The advantage of using tubular tissue is minimized losses due to better current distribution. Another advantage of using tubular tissue in the head coil hypothesis is good ventilation as well as minimal visual disturbances for patients and operators due to the large grooves that can appear in the form of coils between adjacent conductive elements due to the thickness of the tubular tissue. Do it. The conductive loop element can be hypothesized from a double-sided printed circuit board using Teflon resin as the substrate material, which serves as a capacitor of copper alloy, for example. The capacitor sheet metal is composed of overlapping conductive copper regions on opposite sides of the substrate. It may be desirable to form a plurality of discrete pads in one of the capacitor sheet metals that may be indispensable to increase or decrease the capacitance by increasing or decreasing the area of overlap between the sheet metals. I do not know. Many such capacitors are electrically connected to each other in a loop to form the electrically conductive loop element. As an example, to resonate at 62.5 MHz, the capacitor for the head coil has a value of about 95 pF.

The body coil may be hypothesized in a manner similar to that described for the head coil. The shaft's conductive element has an outer diameter of about 0.51 m and is assembled using 38 mm wide copper tape in the form of a cylindrical fiber glass coil about 0.56 m long. The capacitive element is selected at approximately 105 pF. Another body coil made from a copper tube with an outer diameter of 6 mm is used for the conductive element of the shaft.

Additional details of the coil hypothesis can be found by reference to previously related patent applications incorporated herein by reference.

In comparison, the coil of the present invention (ladder network coil) generates a significantly more uniform field than the saddle coil. This is due to the current distribution in the coil of the invention when excited in the preferred mode, especially because it approaches closer to the ideal surface current distribution which gives a completely uniform lateral RF field.

As before, although the particular embodiment has been shown as having eight such axes of conductive elements, fewer than eight or more numbers are advantageously used in practicing the present invention. In addition, the coil may be activated by connecting an RF power source, not shown, to one end of the capacitive element at one of the elements of the axis or one end of the capacitive element at one of the conductive loops. It must be recognized. Other excitation techniques previously described are also applicable to the band pass coil. In addition, each conductive portion of the conductive loop element or the conductive element of the axis, although not shown, is associated with its inductance, which is necessary to achieve the exact angle phase shift in order to generate the desired resonance state.

In the band pass RF coil, the resonance mode is more compressed than in the high pass or low pass coil. The mode is straight between the low frequency limit where the axis and loop impedance is predominantly capacitive and the high frequency limit where the axis and loop impedance is predominantly conducting. In the frequency band in which the mode is present, the capacitor value causes the loop conductor to be induced while the device on the axis is capacitive or vice versa. The first case is similar to a low pass coil in that it is the lowest frequency mode of the set that produces a roughly uniform lateral RF magnetic field. The capacitance is such that when the element of the axis is induced while the element of the circumference is capacitive, then the highest frequency transverse mode is that which produces a uniform transverse field and its arrangement is similar to a high pass RF coil.

Bandpass coil arrangements may be useful to reduce the effects of idle capacitance coupled to or around the image former or to extend slow wave delay line network coils for higher frequencies. For example, capacitance using a high pass coil as the basic arrangement can be inserted into the device of the axis. The capacitor of the axis can be made to erase the capacitance of the device of any axis such that the total impedance of the device of the axis is reduced. Then, the voltage along the element of the axis for a given current will be less than it does not have a capacitor, and the current through the idle capacitance will be less whereby the reduction loss and the harmonic effect are related to the idle capacitance.

From the foregoing, a series of embodiments of coils consistent with the present invention provide current and tuning capacitances distributed in many directions. The NMR RF coil of the present invention also provides a significant improvement in signal sensitivity and B 1 uniformity. The geometry of the coil also allows for improvements in signal-to-noise ratio and cyclically polar excitation.

The present invention has been described with reference to specific embodiments and examples, and other variations and changes may be made by those skilled in the art. Thus, it goes without saying that, within the scope of the claims of the present invention, it may be implemented in other ways than specifically described.

The present invention, as described above, can be utilized as a magnetic field detection or transmission method using a magnetic resonator and the device, the magnetic field detection device of the present invention is a simple wire form housed in a mold made of a non-magnetic material By incorporating a coil type magnetic resonator which is inductively coupled to the loop antenna by being enclosed in a mold made of a loop antenna and a non-magnetic material, the structure is simpler than a conventional MR signal detection device, and can be manufactured with high sensitivity and low price. The features that can be used are also available.

The magnetic resonator according to the present invention, as illustrated in FIG. 2D as a nonmagnetic material, forms an annular metal pattern up and down on a cylindrical printed circuit board (PCB) substrate, and arranges a capacitor therebetween. It may be in one configuration.

The magnetic resonator may include a replaceable capacitor or a variable capacitor, and is configured to control the resonance frequency by exchanging the replaceable capacitor or adjusting the capacitance of the variable capacitor.

In this way, the resonance is generated by the time-varying magnetic field in the magnetic resonator of the coil type inductively coupled to the loop antenna.

Thus, by arranging a plurality of capacitors to form a magnetic resonator in the simple upper and lower loop antenna type metal bands 23 and 25, it is possible to provide a magnetic resonator having high sensitivity and simple structure and low cost. Will be.

In small objects such as small animals, MR magnetic resonance imaging (MRI) apparatus detects a magnetic field of a specific frequency emitted by a hydrogen nucleus contained in a human body and converts it into a two-dimensional or three-dimensional image. It is a device that allows you to examine the internal tissue type. It uses the magnetic field, which is harmless to the human body, compared to computed tomography (CT) or X-ray. It will be able to perform advanced, non-destructive and non-radioactive inspections that are harmless and have excellent contrast and resolution.

In addition, since the present invention can be further miniaturized, small field of view (FOV), high resolution image, and signal-to-noise ratio (SNR) improved MR images, which have not been previously obtained in clinical MRI equipment, can be obtained. May be. Small sized RF-coils for small animals increase filling-factor compared to clinical RF-coils, and smaller size increases overall resistance (R) to improve SNR.Marketability and commercialization prospects Can be sold as lab equipment1. It is possible to study small animals using various size animal models.

In addition, according to the invention of RF-coil for MRI (Magnetic Resonance Imaging) suitable for small animals (Rat, Mouse), it is possible to acquire the image according to the organic, so that it is possible to easily check the specific site targeting using a substance such as contrast agent will be.

Preferred embodiments of the present invention described above are disclosed for purposes of illustration, and those skilled in the art having various ordinary knowledge of the present invention may make various modifications, changes, and additions within the spirit and scope of the present invention. And additions should be considered to be within the scope of the following claims.

Meanwhile, in the detailed description of the present invention, specific embodiments have been described, but it will be apparent to those skilled in the art that various modifications are possible without departing from the scope of the present invention.

In the above description with reference to the accompanying drawings, the present invention has been described with reference to specific shapes and directions, but the present invention may be variously modified and changed by those skilled in the art, and such modifications and changes are included in the scope of the present invention. Should be interpreted as

Claims (3)

The high frequency resonator includes a pair of conductive loop elements at regular intervals along a common vertical axis, each of which includes a plurality of series connected capacitive elements at regular intervals along the circumference thereof. And a plurality of shaft conductive portions electrically connecting the conductive loop elements to each other at points between adjacent capacitive elements of the series connected capacitive elements.
The method of claim 1,
Of the conductive loop with a first RF power source oscillating at a frequency of the resonance mode desired to produce a current distribution that is to be changed according to through the conductive portion of the axis and proportional to the current in the conductive loop element. And a first method for exciting the RF coil across the first input capacitive element in one, where K is an integer between 1 and the total number of coil cuts, N, where K = 1 in the cut with the input capacitor. The animal high frequency resonator characterized by starting with.
The method of claim 1,
The RF coil across the second input capacitive element separated by 90 Hz along the one circumference of the conductive loop element from the first input capacitive element having a second RF power source at the desired resonance mode frequency And a second method for exciting the power source, wherein the power source has a phase difference of 90 GHz with respect to each other.

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