CA1230922A - Nuclear magnetic resonance radio frequency antenna - Google Patents
Nuclear magnetic resonance radio frequency antennaInfo
- Publication number
- CA1230922A CA1230922A CA000482454A CA482454A CA1230922A CA 1230922 A CA1230922 A CA 1230922A CA 000482454 A CA000482454 A CA 000482454A CA 482454 A CA482454 A CA 482454A CA 1230922 A CA1230922 A CA 1230922A
- Authority
- CA
- Canada
- Prior art keywords
- coil
- conductors
- strips
- conductor
- wing
- Prior art date
- Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
- Expired
Links
Classifications
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
- G01R33/00—Arrangements or instruments for measuring magnetic variables
- G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
- G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
- G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
- G01R33/34—Constructional details, e.g. resonators, specially adapted to MR
- G01R33/34007—Manufacture of RF coils, e.g. using printed circuit board technology; additional hardware for providing mechanical support to the RF coil assembly or to part thereof, e.g. a support for moving the coil assembly relative to the remainder of the MR system
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
- G01R33/00—Arrangements or instruments for measuring magnetic variables
- G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
- G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
- G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
- G01R33/34—Constructional details, e.g. resonators, specially adapted to MR
- G01R33/34046—Volume type coils, e.g. bird-cage coils; Quadrature bird-cage coils; Circularly polarised coils
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
- G01R33/00—Arrangements or instruments for measuring magnetic variables
- G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
- G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
- G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
- G01R33/36—Electrical details, e.g. matching or coupling of the coil to the receiver
- G01R33/3628—Tuning/matching of the transmit/receive coil
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
- G01R33/00—Arrangements or instruments for measuring magnetic variables
- G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
- G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
- G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
- G01R33/36—Electrical details, e.g. matching or coupling of the coil to the receiver
- G01R33/3642—Mutual coupling or decoupling of multiple coils, e.g. decoupling of a receive coil from a transmission coil, or intentional coupling of RF coils, e.g. for RF magnetic field amplification
- G01R33/3657—Decoupling of multiple RF coils wherein the multiple RF coils do not have the same function in MR, e.g. decoupling of a transmission coil from a receive coil
Abstract
Nuclear Magnetic Resonance Radio Frequency Antenna Abstract A nuclear magnetic resonance radio frequency coil.
The disclosed coil provides high frequency resonance signals for perturbing a magnetic field within the coil.
The coil is impedance matched and tuned with adjustable capacitors. A balanced configuration is achieved with a co-axial cable chosen to phase shift an energization signal coupled to the coil. The preferred coil is a thin metallic foil having a shorting conductor, four wing conductors, and uniquely shaped parallel cross conductors connecting the shorting and wing conductors.
When mounted to a rf transmissive plastic substrate and energized the coil produces a homogenous field within a region of interest the size of a patient head. A semi-circular balanced feedbar arrangement is used to minimize undesired field contributions.
The disclosed coil provides high frequency resonance signals for perturbing a magnetic field within the coil.
The coil is impedance matched and tuned with adjustable capacitors. A balanced configuration is achieved with a co-axial cable chosen to phase shift an energization signal coupled to the coil. The preferred coil is a thin metallic foil having a shorting conductor, four wing conductors, and uniquely shaped parallel cross conductors connecting the shorting and wing conductors.
When mounted to a rf transmissive plastic substrate and energized the coil produces a homogenous field within a region of interest the size of a patient head. A semi-circular balanced feedbar arrangement is used to minimize undesired field contributions.
Description
~L~3~2 9-a74 Description Nuclezr Ma~netic Resonance Radio Frequenc~_A_tenna 5 Technical Field The present invention relates to nuclear magnetic resonance imaging and more particularly to an im?roveà
resonator for applying radio frequency pulses and receiv-ing low level RF signals over a region of interest.
Backa~ound Art In medical applications, nuclear magnetic reasona-._e ~NY~ can indicate variations in the distribution of ato~ic substances in slices or volumes o~ interest with-in a patient. Such variations can be displayed in a way similar to the distributions provided by a compu~er-ized tomography system. In a nuclear magnetic resonanceexamination, magnetic and rf fields rather than x-radi2.ion scan the body. Resonances caused by these fields are detected as induced signals in one or more detector coil systems. The outputs from these coils are then stor-d and 2nalyzed so that NMR distributions can be dis?laved.
Techniques for producing these im2ges are well known in the art and disclosed in various printed pu~
cations 2nd U.S. patents. Several proposals for app2~z-tus to utilize these procedures are embodied, for exar~?l~,in U.S. Patent Nos. 4,454,474 to Young, 4,3~4,255 to Young e. al, and 4,379,262 to Younq.
The .echniques disclosed in the above mentioned prior 2rt patents involve selection of a planar slice of interest in the body and application of 2 strong ~agne-ic field gradien. in a direc.ion perpendicular 'o the slice. This field is perturbed in a perpendicular direction in the plane of th-e slice. The direction of the perturbation is continuously varied by a procedure documented in the literature.
,i , .
~23~9;~ -The effect of this perturbation is to introduce a dispersion in nuclear resonance frequenciés which re~urn to their original unperturbed state in ways ch~rac.eris- c of the structure within the slice of interest. Re?e.i~ion of this procedure for different directions can give many signals for each slice of interest which are then used to construct cross-sectional images descriptive Oc the internal structure of the patient slice.
The radio frequency perturbation excites the nucleii by realigning the macroscopic magnetization or masnetic moment wi.hin the cross-section of interest. This radio frequency energization is performed at the Lar~or frequen-cy. This frequency is related to a constant descriptive of the nucleii making up the region o interest an~ the masnetic field gradient imposed during perturbation.
Experience in NMR imaging indicates that the scan-ning times can be reduced and spacial resolution of ~.~R
images can be lncreased by increasing this field to higher levels. Since the Larmor frequency of a given nucleii is directly proportional to the field strength, this increzse in field strength must be accompanied bv highe. ,requencies for RF energization. In the prior art this energization was accomplished with suitcbly designed energization coils which generate perturbation fields and in some instances are also used for detecting signals caused by resonances set up within the region of interest.
Transmission and reception of radio frequency sis-nals for NMR imaging requires a resonant radiating struc-ture, often called an rf coil, meeting severGl cri.e- z.
The resonant point of the structure must be hish enous;.
to a~loh ?roper tur.ing 2t the Crequency OL inteLeSt ~C
the structure must have sufficiently high "Q" to provide good signal to noise perform~nce in the receive mode.
~'~3~2Z
Generally~ in small volume nuclear magnetic reson-ance systems an unbalanced feed system and coil configur-ation is used. A simple reactive element is used as an impedance matching component and a second reactive ele-ment is used in parallel with the coil structure totune the coil to an appropriate frequency.
For large volume nuclear magnetic resonance applica-tions, however, suc,h as a head imaging system a balanced coil system is preferred. This is preferrable since under sample loading the coil system will be less in-fluenced than an asymmetrical system.
As the frequency of operation is raised, however, the effectiveness of a symmetrical matching system is limited by a variet~ of factors. The reactive components become unmanageably small and are also subjected tc extremely high peak voltages. The stray capacitance of the RF coil network eventually makes it impossible to match the network to a useful impedence. In addition to problems in achieving proper energization frequencies, use of larger RF coils creates problems in achieving a unilorm magnetic field over the region to be perturbed.
Various prior art proposals to provide new and imprcved RF energization and detection coils are discus-sed in the literature. A publication entitled "Slotted Tube Resona.or: A new N~R probe head at high observing frequencies" by Schneider and Dullenkopf discusses a resonator for use at high frequencies. This work was the first of a number of similar prior art publications discussing NM~ resonator structures. ~uch of this work, however, has been conducted with extremely small dime~-sional s~ructures which do not encounter the diffic~lties encGuntered ~hen imaging a cross-section of a head.
The task of converting a resonator coil for use in small structure an~lysis into a de~ice suitable for NMR medic21 imasing is not a straight forhard extension of this prior work.
:~3~
.
Disclosure of Invention The invention is particul2rly suited,~or N~ im2sins of a human head. A resonator having a length and dia~.e.e~
of approximately 30 centimeters is constructed ,o provide a homogeneous magnetic field in the region of the head that does not unreasonably degrade with sam?le loading and is easily tuned over a wide frequency range.
~n antenna or coil arrangement constructed in accord-ance with the invention both transmitts and receives high frequency energy in the range of 30 to 95 megaher.z.
The disclosed resonator includes a cylindrical base of a diameter suitable for enclosing the human head and a metallic foil coupled to the base and forming the antenn2 structure with a resonant fre~uency of about 30 to 95 megahertz. The sel~ resonant frequency of the structure is well in excess of 100 megahertz. This resonant struc-ture includes a pair of diametrically opposed arcuate electrical conductors with each conductor subtending 2n arc of between 75 and 85 degrees. Short circuiting conductors interconnect these conductors at one end and win~ strips extend circumferentially from each of the cons~-ctors at the other end of the resonator. Enerqiza-tion signa1s are applied to the resonator through conduc-tive feed strips which interconnect the wing stri?s.
The arr2ngement between feed strips, wing strips and conductor strips causes uniform magnetic fielc ~ithi~
the region of interest, i.e. the head. In a ~referred embodiment, the conductor strips are circumferentiail;
spaced parallel conductor strips with each strip increcs-ing in cross-sectional area from its center towards each end. This configuration of the conductor s.-l?s enh~.~ces the u"i~or~ity Oc the magnetic fielc whil- ^c-affecting the tunability of the structure.
An interface circuit is preferably coupled between ~5 the disclosed resonator and a standard 50 ohm input 1, .
~z3C~92;Z
cable. The resonator also acts as a pick up coil so that resonances within a patient slice of~ nterest induce electrical signals which are detected, amplified, and utilized in constructing an NMR image- To achieve pro~er impedance matching and resonance the disclosed resonator is coupled to the 50 ohm input cable through three adjus~-able capacitors and a half wavelength balun coaxi21 cable -The resonatOr is a quarter wavelength antenna which can be easily tuned and matched in a transmit mode ofoperation and effectively coupled to a pre-amplifier for generating output signals for use in NM~ imaging.
From the above it should be appreciated that one object of the invention is a antenna structure suitable for r sigr.al generation and reception at high fre~uen~ies with a geometry large enough for head imaging. Other objects, advantages and features of the invention will becol"e better understood when a detailed description of a preferred embodiment of the invention is described in conjunction with the accompanying drawings.
Brief DescriPtion of the Drawinqs ~ igure 1 is a perspective view of an N~R imaging station.
Pisure 2 is a perspective view of a resona,or pro~e for providing rf signals in the vicinity of a patien,'s head.
Figure 3 is a top plan view of a foil configured to form the Figure 2 resonator.
Figure 4 is a schematic circuit diagram o, the resonator showing use of three adjustzble cap2citor for tuning and impedance ~atching.
~ igure 5 is a schema.ic snowing an entire ~ ? ~
coil system for both transmitting and receiving resonance signals from within a region of interest.
.
~3(3~
Figure 6 shows a filter for reducing transmitter noise.
Figures 7-9 are graphical representations showins magnetic field uniformity in the region encircled by the antenna.
Best Mode for Carryinq out the Invention Turning now to the drawings and in particular Figur~
1, an imaging station for an N~lR scanner 10 is disclosed.
The scanner 10 includes a large encircling magnet 12 for generating magnetic fields of between 1.5 and 2 Tesla within a patient aperture 14. Shown positioned in proximity to the magnet 12 is a patient couch 16 having a headrest 18. The patient is positioned on the couch in a prone position and then moved into the patient lS aperture 14 for N~R scannin~.
During a head scan a probe coil or resonator 20 i5 moved on rollers 22 so that the patient's head is posi-tioned within the coil 20. In accordance with ~echnicues well known in the nuclear magnetic resonance imaging art, the magnet 14 is energized to produce a strong magnetic field having a gradient to selectively choose
resonator for applying radio frequency pulses and receiv-ing low level RF signals over a region of interest.
Backa~ound Art In medical applications, nuclear magnetic reasona-._e ~NY~ can indicate variations in the distribution of ato~ic substances in slices or volumes o~ interest with-in a patient. Such variations can be displayed in a way similar to the distributions provided by a compu~er-ized tomography system. In a nuclear magnetic resonanceexamination, magnetic and rf fields rather than x-radi2.ion scan the body. Resonances caused by these fields are detected as induced signals in one or more detector coil systems. The outputs from these coils are then stor-d and 2nalyzed so that NMR distributions can be dis?laved.
Techniques for producing these im2ges are well known in the art and disclosed in various printed pu~
cations 2nd U.S. patents. Several proposals for app2~z-tus to utilize these procedures are embodied, for exar~?l~,in U.S. Patent Nos. 4,454,474 to Young, 4,3~4,255 to Young e. al, and 4,379,262 to Younq.
The .echniques disclosed in the above mentioned prior 2rt patents involve selection of a planar slice of interest in the body and application of 2 strong ~agne-ic field gradien. in a direc.ion perpendicular 'o the slice. This field is perturbed in a perpendicular direction in the plane of th-e slice. The direction of the perturbation is continuously varied by a procedure documented in the literature.
,i , .
~23~9;~ -The effect of this perturbation is to introduce a dispersion in nuclear resonance frequenciés which re~urn to their original unperturbed state in ways ch~rac.eris- c of the structure within the slice of interest. Re?e.i~ion of this procedure for different directions can give many signals for each slice of interest which are then used to construct cross-sectional images descriptive Oc the internal structure of the patient slice.
The radio frequency perturbation excites the nucleii by realigning the macroscopic magnetization or masnetic moment wi.hin the cross-section of interest. This radio frequency energization is performed at the Lar~or frequen-cy. This frequency is related to a constant descriptive of the nucleii making up the region o interest an~ the masnetic field gradient imposed during perturbation.
Experience in NMR imaging indicates that the scan-ning times can be reduced and spacial resolution of ~.~R
images can be lncreased by increasing this field to higher levels. Since the Larmor frequency of a given nucleii is directly proportional to the field strength, this increzse in field strength must be accompanied bv highe. ,requencies for RF energization. In the prior art this energization was accomplished with suitcbly designed energization coils which generate perturbation fields and in some instances are also used for detecting signals caused by resonances set up within the region of interest.
Transmission and reception of radio frequency sis-nals for NMR imaging requires a resonant radiating struc-ture, often called an rf coil, meeting severGl cri.e- z.
The resonant point of the structure must be hish enous;.
to a~loh ?roper tur.ing 2t the Crequency OL inteLeSt ~C
the structure must have sufficiently high "Q" to provide good signal to noise perform~nce in the receive mode.
~'~3~2Z
Generally~ in small volume nuclear magnetic reson-ance systems an unbalanced feed system and coil configur-ation is used. A simple reactive element is used as an impedance matching component and a second reactive ele-ment is used in parallel with the coil structure totune the coil to an appropriate frequency.
For large volume nuclear magnetic resonance applica-tions, however, suc,h as a head imaging system a balanced coil system is preferred. This is preferrable since under sample loading the coil system will be less in-fluenced than an asymmetrical system.
As the frequency of operation is raised, however, the effectiveness of a symmetrical matching system is limited by a variet~ of factors. The reactive components become unmanageably small and are also subjected tc extremely high peak voltages. The stray capacitance of the RF coil network eventually makes it impossible to match the network to a useful impedence. In addition to problems in achieving proper energization frequencies, use of larger RF coils creates problems in achieving a unilorm magnetic field over the region to be perturbed.
Various prior art proposals to provide new and imprcved RF energization and detection coils are discus-sed in the literature. A publication entitled "Slotted Tube Resona.or: A new N~R probe head at high observing frequencies" by Schneider and Dullenkopf discusses a resonator for use at high frequencies. This work was the first of a number of similar prior art publications discussing NM~ resonator structures. ~uch of this work, however, has been conducted with extremely small dime~-sional s~ructures which do not encounter the diffic~lties encGuntered ~hen imaging a cross-section of a head.
The task of converting a resonator coil for use in small structure an~lysis into a de~ice suitable for NMR medic21 imasing is not a straight forhard extension of this prior work.
:~3~
.
Disclosure of Invention The invention is particul2rly suited,~or N~ im2sins of a human head. A resonator having a length and dia~.e.e~
of approximately 30 centimeters is constructed ,o provide a homogeneous magnetic field in the region of the head that does not unreasonably degrade with sam?le loading and is easily tuned over a wide frequency range.
~n antenna or coil arrangement constructed in accord-ance with the invention both transmitts and receives high frequency energy in the range of 30 to 95 megaher.z.
The disclosed resonator includes a cylindrical base of a diameter suitable for enclosing the human head and a metallic foil coupled to the base and forming the antenn2 structure with a resonant fre~uency of about 30 to 95 megahertz. The sel~ resonant frequency of the structure is well in excess of 100 megahertz. This resonant struc-ture includes a pair of diametrically opposed arcuate electrical conductors with each conductor subtending 2n arc of between 75 and 85 degrees. Short circuiting conductors interconnect these conductors at one end and win~ strips extend circumferentially from each of the cons~-ctors at the other end of the resonator. Enerqiza-tion signa1s are applied to the resonator through conduc-tive feed strips which interconnect the wing stri?s.
The arr2ngement between feed strips, wing strips and conductor strips causes uniform magnetic fielc ~ithi~
the region of interest, i.e. the head. In a ~referred embodiment, the conductor strips are circumferentiail;
spaced parallel conductor strips with each strip increcs-ing in cross-sectional area from its center towards each end. This configuration of the conductor s.-l?s enh~.~ces the u"i~or~ity Oc the magnetic fielc whil- ^c-affecting the tunability of the structure.
An interface circuit is preferably coupled between ~5 the disclosed resonator and a standard 50 ohm input 1, .
~z3C~92;Z
cable. The resonator also acts as a pick up coil so that resonances within a patient slice of~ nterest induce electrical signals which are detected, amplified, and utilized in constructing an NMR image- To achieve pro~er impedance matching and resonance the disclosed resonator is coupled to the 50 ohm input cable through three adjus~-able capacitors and a half wavelength balun coaxi21 cable -The resonatOr is a quarter wavelength antenna which can be easily tuned and matched in a transmit mode ofoperation and effectively coupled to a pre-amplifier for generating output signals for use in NM~ imaging.
From the above it should be appreciated that one object of the invention is a antenna structure suitable for r sigr.al generation and reception at high fre~uen~ies with a geometry large enough for head imaging. Other objects, advantages and features of the invention will becol"e better understood when a detailed description of a preferred embodiment of the invention is described in conjunction with the accompanying drawings.
Brief DescriPtion of the Drawinqs ~ igure 1 is a perspective view of an N~R imaging station.
Pisure 2 is a perspective view of a resona,or pro~e for providing rf signals in the vicinity of a patien,'s head.
Figure 3 is a top plan view of a foil configured to form the Figure 2 resonator.
Figure 4 is a schematic circuit diagram o, the resonator showing use of three adjustzble cap2citor for tuning and impedance ~atching.
~ igure 5 is a schema.ic snowing an entire ~ ? ~
coil system for both transmitting and receiving resonance signals from within a region of interest.
.
~3(3~
Figure 6 shows a filter for reducing transmitter noise.
Figures 7-9 are graphical representations showins magnetic field uniformity in the region encircled by the antenna.
Best Mode for Carryinq out the Invention Turning now to the drawings and in particular Figur~
1, an imaging station for an N~lR scanner 10 is disclosed.
The scanner 10 includes a large encircling magnet 12 for generating magnetic fields of between 1.5 and 2 Tesla within a patient aperture 14. Shown positioned in proximity to the magnet 12 is a patient couch 16 having a headrest 18. The patient is positioned on the couch in a prone position and then moved into the patient lS aperture 14 for N~R scannin~.
During a head scan a probe coil or resonator 20 i5 moved on rollers 22 so that the patient's head is posi-tioned within the coil 20. In accordance with ~echnicues well known in the nuclear magnetic resonance imaging art, the magnet 14 is energized to produce a strong magnetic field having a gradient to selectively choose
2 slice or region of patient interest. With the probe coil 20 encircling the patient's head, the coil is ener-gized with z high frequency (between 30 and 95 megaher~z~
signal which sets up a time varying magnetic field in the region of interest. Various techniques are known within the art for pulsing the probe coil in ways to produce meaningful resonance information which can be utilized in ~MR imaging. The particular configuratior.
of the disclosed coil 20 allows high frequency enersizz-tion necessary to cause resonance of the spin sys e, z the high rzgne~ic fields generated by ~he magne~ 12.
At such high frequencies, the disclosed pro~e 20 proouces uniform magnetic fields which do no, exhibit undue Q
degradation with sample loading.
~23~ 2 Turning now to Figures 2-4, details of the construc-tion of the probe coil are discussed. A cylindrical base 30 formed from an acrylic material forms a surface to which a metallic foil can be affixed. The base 30 has phy~ical dimensions such that a patient's head can be inserted within the base and the probe coil 20 ener-gized in conjunction with generation of the high strength magnetic field. Two copper foil resonator sections 32 having a thickness of .0635 millimeter are affixed to an outer surface of the base 30 in a configuration shown in rigure 2. The foils are self adhesive with a backin~
laye- which is removed prior to application to the base 30. One of the foil sections 32 is shown in plan view prior to mounting to the substrate 30 in Figure 3. The 1~ physlcal dimensions of this foil are shown in that figu~e.
The thickness of the foil is chosen to be approxi-mate1y seven skin depths at the resonant frequency.
Use of this thickness causes the resonator 20 to be essentially transparent to the high strength field grad-ients generated by the magnet 12. This minimizes the $
generation of eddy currents within the foil by this high s.rength magnetic field gradient which would be undesirable since the induced eddy currents would produce their cwn magnetic field in addition to the desired 2S homoseneous rf field.
Each foil segment 32 includes a shorting strip 34 and a wins strip 36. These two strips 34, 36 are inter-connected by conductor strips 38 which are parAllel to each other and nonuniform in width along their length.
Preferably these conductor strips 38 are narrow in .he middle and widen as they 2pproach the shorting s~rip 3~ 7 and wing strip 36. As seen mos~ clearly in Figure 2, when zffixed to the outside surface of the substrate 30, the two foil sesments 32 contact each other at the ends of the shorting strips 34 and define a 1 cm ga?
between the ends of the wing strips 36.
~Z3~922 An end portion 36a of each wing 36 i ~connected by a feed bar 40 having a midpoint 42 connected to an inter-face circuit 110 (Figure 5) by copper strips (not sho~n3.
The feed bars both energize the probe and transmit reso-nance signals generated from within the patient regionof interest. The feed bars 40 form a semicircle 2 cm wide each of the same thickness as the foil and are mounted to an acrylic substrate.
The resonator 20 is energized with a high frequen~y output from a transmitter 112. A preferred transmitter is available from Amplifier Research under Model No. 2000 .~L8 and produces an alternating current voltaqe a few hundred volts in magnitude In order to interface the resonator 20 to a standard S0 ohm unbalanced transrnission line 11~ a half wavelength balun 115 (Figures 4 and 5) is utiliæed. The balun 115 is constructed from 50 ohm coaxial cable, with a velocity factor of 0.80 or 0.66.
The total length of the balun is 1.875 meter using a velocity factor of 0.80 at 64 megahertz. Since the signal traveling in this cable is delayed by one half ~avelength the phase of the voltage at one end of the czble is 180 shifted from the other end. Thus, the voltcges at each end of the cable are equal in amplituae and opposite in phase. The current at an input node 116 divides eaually between the load and the balun phas-ing line. Thus, the resonator matching network sees a voltage double that of the input voltage, and a current e~ual to half the applied current. This causes the impedance a~ the output of the m~tching ne~work to be ~our times the input line impedance.
A m2tching network 120 having Lhree adjustable capaci-toLs 122, 124, 126 is used to tune the resona OL
20 and imped2nce match the high-impedance resonator 20 ~ith the 200 ohm balanced input. Model Number CACA 125 ~5 vacuum capacitors from ITT Jennings of San Jose, Californi2 ~L~2,3(~9;~z are preferred. Representative values of these capacitors are 60 pico farads for the parallel capacitor 122 and 12 picofarads for the two series capacitors 124, 126.
These values are representative and are tuned to optimize S performance of the resonator 20.
To utili~e the resonator 20 as ~oth a transmitter and receiver, a multiplex circuit 130 (Figure 5) couples the resonator and balancing network to both the transmitter 112 and a pre-ami?lifier 132. ~he multiplex network includes a plurality of dio~es 134 and two quarter wave-length cables 136, 138.
In the transmit mode the large magnitude signals from the transmitter 112 forward bias the diodes 134.
The quarter wavelength cable 136 consu~Tes no net power since the cable inverts the terminating impedance and no signal from the transmitter reaches the pre-amplifier 132.
In a receive mode the goal is to couple induced signals in the resonator 20 to the pre-amplifier 132.
These signals see a half wavelength cable since the two qua_te wave cables 136, 138 act as a single half wave cable.
A noise filter circuit 140 (Figure 6) couples the transmitter 112 to the multiplexer circuit 130 and in-cluaes a ?lurality of diodes 142 and two quarter wave~
length cables 144, 146 which function in a way simi~ar to the diodes 13~ and cables 136, 138 of the multiplex circuit 130. In the transmit mode the diodes are for-ward biased, shorting the cable 146 to a quarter wave-length cable 144. No net pc~wer is consumed by the cable 144. In a receive mode the ~wo czbles 144, 146 act as a singie half wavelength cable. Any noise from the transmitter is blocked since the half waveleng.h cable presents a virtual short.
~L~30922 The pre-amplifier 132 is coupled to other ap~ar~tus known in the nuclear magnetic resonating art for conver--ing signals from the resonator into si9nals s~itable for imaginq. The resonator 20 has an unloaded "Q" of about 300 and a loaded "Qn of approximately 50. A ver~
good match to the 50 ohm transmission line is achieved with reflected power levels under two percent.
Field uniformity is presented in Figures 7-9 where a plot of variations and magnetic field strength ~ith position in the X, Y, and Z directions as defined in Figure 1 are disclosed. The origin of this co-ordinate axis is a point centered within the resonator 20 halfway between the shorting and wing conductors. The data presented in Figures 7-9 was generated with a probe coil energization of 64.5 megahertz t an unloaded "Q" of 260 and a loaded "Q" of 55. The X and Y uniformit~ in field is excellent and by properly positioning the reson-ator 20 along the Z axis uniformity within a region of interest as defined by the field gradient of the magnet 12 can be aGhieved.
The disclosed design fulfills all the requiremen.s for high quality head imaging at field strengths of l.S
Tesla. The operating parameters of the resonator 20, however, should not be viewed as limiting the invention and field strengths of 2.0 Tesla and resonance frequen- !
cies of 85 megahertz are possible. It is the intent that the invention cover all modifications and/or zlter-ations following within the spirit or scope of the zp?e~-ded claims. ;
signal which sets up a time varying magnetic field in the region of interest. Various techniques are known within the art for pulsing the probe coil in ways to produce meaningful resonance information which can be utilized in ~MR imaging. The particular configuratior.
of the disclosed coil 20 allows high frequency enersizz-tion necessary to cause resonance of the spin sys e, z the high rzgne~ic fields generated by ~he magne~ 12.
At such high frequencies, the disclosed pro~e 20 proouces uniform magnetic fields which do no, exhibit undue Q
degradation with sample loading.
~23~ 2 Turning now to Figures 2-4, details of the construc-tion of the probe coil are discussed. A cylindrical base 30 formed from an acrylic material forms a surface to which a metallic foil can be affixed. The base 30 has phy~ical dimensions such that a patient's head can be inserted within the base and the probe coil 20 ener-gized in conjunction with generation of the high strength magnetic field. Two copper foil resonator sections 32 having a thickness of .0635 millimeter are affixed to an outer surface of the base 30 in a configuration shown in rigure 2. The foils are self adhesive with a backin~
laye- which is removed prior to application to the base 30. One of the foil sections 32 is shown in plan view prior to mounting to the substrate 30 in Figure 3. The 1~ physlcal dimensions of this foil are shown in that figu~e.
The thickness of the foil is chosen to be approxi-mate1y seven skin depths at the resonant frequency.
Use of this thickness causes the resonator 20 to be essentially transparent to the high strength field grad-ients generated by the magnet 12. This minimizes the $
generation of eddy currents within the foil by this high s.rength magnetic field gradient which would be undesirable since the induced eddy currents would produce their cwn magnetic field in addition to the desired 2S homoseneous rf field.
Each foil segment 32 includes a shorting strip 34 and a wins strip 36. These two strips 34, 36 are inter-connected by conductor strips 38 which are parAllel to each other and nonuniform in width along their length.
Preferably these conductor strips 38 are narrow in .he middle and widen as they 2pproach the shorting s~rip 3~ 7 and wing strip 36. As seen mos~ clearly in Figure 2, when zffixed to the outside surface of the substrate 30, the two foil sesments 32 contact each other at the ends of the shorting strips 34 and define a 1 cm ga?
between the ends of the wing strips 36.
~Z3~922 An end portion 36a of each wing 36 i ~connected by a feed bar 40 having a midpoint 42 connected to an inter-face circuit 110 (Figure 5) by copper strips (not sho~n3.
The feed bars both energize the probe and transmit reso-nance signals generated from within the patient regionof interest. The feed bars 40 form a semicircle 2 cm wide each of the same thickness as the foil and are mounted to an acrylic substrate.
The resonator 20 is energized with a high frequen~y output from a transmitter 112. A preferred transmitter is available from Amplifier Research under Model No. 2000 .~L8 and produces an alternating current voltaqe a few hundred volts in magnitude In order to interface the resonator 20 to a standard S0 ohm unbalanced transrnission line 11~ a half wavelength balun 115 (Figures 4 and 5) is utiliæed. The balun 115 is constructed from 50 ohm coaxial cable, with a velocity factor of 0.80 or 0.66.
The total length of the balun is 1.875 meter using a velocity factor of 0.80 at 64 megahertz. Since the signal traveling in this cable is delayed by one half ~avelength the phase of the voltage at one end of the czble is 180 shifted from the other end. Thus, the voltcges at each end of the cable are equal in amplituae and opposite in phase. The current at an input node 116 divides eaually between the load and the balun phas-ing line. Thus, the resonator matching network sees a voltage double that of the input voltage, and a current e~ual to half the applied current. This causes the impedance a~ the output of the m~tching ne~work to be ~our times the input line impedance.
A m2tching network 120 having Lhree adjustable capaci-toLs 122, 124, 126 is used to tune the resona OL
20 and imped2nce match the high-impedance resonator 20 ~ith the 200 ohm balanced input. Model Number CACA 125 ~5 vacuum capacitors from ITT Jennings of San Jose, Californi2 ~L~2,3(~9;~z are preferred. Representative values of these capacitors are 60 pico farads for the parallel capacitor 122 and 12 picofarads for the two series capacitors 124, 126.
These values are representative and are tuned to optimize S performance of the resonator 20.
To utili~e the resonator 20 as ~oth a transmitter and receiver, a multiplex circuit 130 (Figure 5) couples the resonator and balancing network to both the transmitter 112 and a pre-ami?lifier 132. ~he multiplex network includes a plurality of dio~es 134 and two quarter wave-length cables 136, 138.
In the transmit mode the large magnitude signals from the transmitter 112 forward bias the diodes 134.
The quarter wavelength cable 136 consu~Tes no net power since the cable inverts the terminating impedance and no signal from the transmitter reaches the pre-amplifier 132.
In a receive mode the goal is to couple induced signals in the resonator 20 to the pre-amplifier 132.
These signals see a half wavelength cable since the two qua_te wave cables 136, 138 act as a single half wave cable.
A noise filter circuit 140 (Figure 6) couples the transmitter 112 to the multiplexer circuit 130 and in-cluaes a ?lurality of diodes 142 and two quarter wave~
length cables 144, 146 which function in a way simi~ar to the diodes 13~ and cables 136, 138 of the multiplex circuit 130. In the transmit mode the diodes are for-ward biased, shorting the cable 146 to a quarter wave-length cable 144. No net pc~wer is consumed by the cable 144. In a receive mode the ~wo czbles 144, 146 act as a singie half wavelength cable. Any noise from the transmitter is blocked since the half waveleng.h cable presents a virtual short.
~L~30922 The pre-amplifier 132 is coupled to other ap~ar~tus known in the nuclear magnetic resonating art for conver--ing signals from the resonator into si9nals s~itable for imaginq. The resonator 20 has an unloaded "Q" of about 300 and a loaded "Qn of approximately 50. A ver~
good match to the 50 ohm transmission line is achieved with reflected power levels under two percent.
Field uniformity is presented in Figures 7-9 where a plot of variations and magnetic field strength ~ith position in the X, Y, and Z directions as defined in Figure 1 are disclosed. The origin of this co-ordinate axis is a point centered within the resonator 20 halfway between the shorting and wing conductors. The data presented in Figures 7-9 was generated with a probe coil energization of 64.5 megahertz t an unloaded "Q" of 260 and a loaded "Q" of 55. The X and Y uniformit~ in field is excellent and by properly positioning the reson-ator 20 along the Z axis uniformity within a region of interest as defined by the field gradient of the magnet 12 can be aGhieved.
The disclosed design fulfills all the requiremen.s for high quality head imaging at field strengths of l.S
Tesla. The operating parameters of the resonator 20, however, should not be viewed as limiting the invention and field strengths of 2.0 Tesla and resonance frequen- !
cies of 85 megahertz are possible. It is the intent that the invention cover all modifications and/or zlter-ations following within the spirit or scope of the zp?e~-ded claims. ;
Claims (4)
PROPERTY OR PRIVILEGE IS CLAIMED ARE DEFINED AS FOLLOWS:
1. An antenna arrangement for transmitting and receiving high frequency energy in the range of 30 to 95 MHZ for use in NMR patient imaging, comprising an insulat-ing cylindrical base of diameter suitable for enclosing a human head, metallic foil on said base forming an anten-na structure having a resonant frequency in the range of 30 to 95 MHZ and including a pair of diametrically oppo-site arcuate electrical conductors, each conductor sub-tending a predetermined arc, short circuiting strips interconnecting said conductors at one end thereof, wing strips extending circumferentially from each side of each conductor at the other end -thereof, and conductive feed strips interconnecting the wing strips of each conductor and lying in planes transverse to the planes of the wing strips.
2. An antenna arrangement as claimed in Claim 1, wherein each of said conductors is configured to en-hance the magnetic field generated within the cylinder during use for imaging purposes.
3. An antenna arrangement as claimed in Claim 2, wherein each of said conductors is shaped to provide a plurality of circumferentially spaced parallel conductor strips, each strip increasing in cross-sectional area from the centre of its length toward each end.
4. The antenna of Claim 1, wherein each conduc-tor subtends 75°-85° of arc.
Priority Applications (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA000547675A CA1236880A (en) | 1984-08-16 | 1987-09-23 | Nuclear magnetic resonance radio frequency antenna |
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| US06/641,570 US4634980A (en) | 1984-08-16 | 1984-08-16 | Nuclear magnetic resonance radio frequency antenna |
| US641,570 | 1984-08-16 |
Related Child Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA000547675A Division CA1236880A (en) | 1984-08-16 | 1987-09-23 | Nuclear magnetic resonance radio frequency antenna |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA1230922A true CA1230922A (en) | 1987-12-29 |
Family
ID=24572945
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA000482454A Expired CA1230922A (en) | 1984-08-16 | 1985-05-27 | Nuclear magnetic resonance radio frequency antenna |
Country Status (5)
| Country | Link |
|---|---|
| US (1) | US4634980A (en) |
| EP (1) | EP0171972B1 (en) |
| JP (1) | JPS6162454A (en) |
| CA (1) | CA1230922A (en) |
| DE (1) | DE3580949D1 (en) |
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| US4755756A (en) * | 1986-02-18 | 1988-07-05 | Mitsubishi Denki Kabushiki Kaisha | Radio frequency coil for nuclear magnetic resonance imaging |
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| US4791371A (en) * | 1986-11-17 | 1988-12-13 | Memorial Hospital For Cancer And Allied Diseases | Apparatus useful in magnetic resonance imaging |
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| IL82658A (en) * | 1987-05-26 | 1990-12-23 | Elscint Ltd | Balun circuit for radio frequency coils in mr systems |
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| NL8802609A (en) * | 1988-10-24 | 1990-05-16 | Philips Nv | MAGNETIC RESONANCE DEVICE WITH OPTIMIZED DETECTION FIELD. |
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| DE4024599C2 (en) * | 1989-08-16 | 1996-08-14 | Siemens Ag | High-frequency antenna of an MRI scanner |
| US5075624A (en) * | 1990-05-29 | 1991-12-24 | North American Philips Corporation | Radio frequency quadrature coil construction for magnetic resonance imaging (mri) apparatus |
| US5250901A (en) * | 1991-11-07 | 1993-10-05 | The Regents Of The University Of California | Open architecture iron core electromagnet for MRI using superconductive winding |
| FI91357C (en) * | 1991-11-15 | 1994-06-27 | Picker Nordstar Oy | Anatomical support for an MRI device |
| US5379767A (en) * | 1992-09-02 | 1995-01-10 | The Regents Of The University Of California | MRI RF coil using zero-pitch solenoidal winding |
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| US5515855A (en) * | 1994-08-05 | 1996-05-14 | Sloan-Kettering Institute For Cancer Research | Dome-shaped resonator for nuclear magnetic resonance imaging and spectroscopy |
| DE19535257A1 (en) * | 1995-09-22 | 1997-03-27 | Philips Patentverwaltung | MR arrangement for determining the nuclear magnetization distribution with a surface coil arrangement |
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| JPH1033494A (en) * | 1996-07-19 | 1998-02-10 | Shimadzu Corp | Magnetic resonance tomography apparatus |
| US6711430B1 (en) | 1998-10-09 | 2004-03-23 | Insight Neuroimaging Systems, Inc. | Method and apparatus for performing neuroimaging |
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| US6798206B2 (en) | 1998-11-25 | 2004-09-28 | Medrad, Inc. | Neurovascular coil system and interface and system therefor and method of operating same in a multitude of modes |
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| US7598739B2 (en) * | 1999-05-21 | 2009-10-06 | Regents Of The University Of Minnesota | Radio frequency gradient, shim and parallel imaging coil |
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| KR20030036663A (en) * | 2000-07-31 | 2003-05-09 | 리전츠 오브 더 유니버스티 오브 미네소타 | Radio frequency magnetic field unit |
| US7084629B2 (en) | 2002-11-27 | 2006-08-01 | Medrad, Inc. | Parallel imaging compatible birdcage resonator |
| CA2565248C (en) * | 2004-05-07 | 2014-07-08 | Regents Of The University Of Minnesota | Multi-current elements for magnetic resonance radio frequency coils |
| DE102006008724A1 (en) * | 2006-02-24 | 2007-11-08 | Siemens Ag | magnetic resonance system |
| US20120133365A1 (en) | 2010-11-26 | 2012-05-31 | Mentis, Llc | System of receive coils and pads for use with magnetic resonance imaging |
| USD979080S1 (en) * | 2019-06-27 | 2023-02-21 | Reza Jalinous | MRI tower coil stand |
| USD1007683S1 (en) * | 2020-09-28 | 2023-12-12 | Siemens Healthcare Gmbh | Medical imaging device |
| USD1001284S1 (en) * | 2020-11-17 | 2023-10-10 | Siemens Healthcare Gmbh | Medical imaging device |
| USD1006230S1 (en) * | 2021-02-26 | 2023-11-28 | Siemens Medical Solutions Usa, Inc. | SPECT/CT imaging scanner |
| USD987081S1 (en) * | 2021-02-26 | 2023-05-23 | Siemens Medical Solutions Usa, Inc. | Spect/CT imaging system |
| USD1005492S1 (en) | 2021-02-26 | 2023-11-21 | Siemens Medical Solutions Usa, Inc. | Patient imaging bed |
| USD998800S1 (en) | 2021-02-26 | 2023-09-12 | Siemens Medical Solutions Usa, Inc. | Patient imaging bed ruler |
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| USD1013175S1 (en) * | 2022-02-15 | 2024-01-30 | Siemens Healthcare Gmbh | Medical imaging device |
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| US4446431A (en) * | 1981-08-24 | 1984-05-01 | Monsanto Company | Double-tuned single coil probe for nuclear magnetic resonance spectrometer |
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-
1984
- 1984-08-16 US US06/641,570 patent/US4634980A/en not_active Expired - Fee Related
-
1985
- 1985-05-27 CA CA000482454A patent/CA1230922A/en not_active Expired
- 1985-07-31 DE DE8585305445T patent/DE3580949D1/en not_active Expired - Lifetime
- 1985-07-31 EP EP85305445A patent/EP0171972B1/en not_active Expired
- 1985-08-16 JP JP60180397A patent/JPS6162454A/en active Granted
Also Published As
| Publication number | Publication date |
|---|---|
| DE3580949D1 (en) | 1991-01-31 |
| US4634980A (en) | 1987-01-06 |
| EP0171972A2 (en) | 1986-02-19 |
| JPH0580210B2 (en) | 1993-11-08 |
| EP0171972B1 (en) | 1990-12-19 |
| JPS6162454A (en) | 1986-03-31 |
| EP0171972A3 (en) | 1987-04-15 |
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