NMR quadrature detection array

A NMR Quadrature Detection Array includes two generally linear arrays; one oriented in a horizontal plane and sensitive to vertically oriented changing magnetic flux, and the other oriented in a vertical plane bisecting the horizontal plane and sensitive to horizontally oriented flux changes. Each horizontal coil is electronically paired with a vertical coil such that their NMR signals are combined in a quadrature fashion before transmission to the NMR system preamplifier. The coil arrays are sized to accommodate easy positioning within the NMR system table top assembly as well as to facilitate rapid and easy patient positioning about the coil elements.

TECHNICAL FIELD 
The present invention relates generally to magnetic resonance imaging (MRI) 
and more particularly to local coils for use in receiving MRI signals. 
BACKGROUND OF THE INVENTION 
A. Magnetic Resonance Imaging 
Magnetic resonance imaging (MRI) refers generally to a form of clinical 
imaging based upon the principles of nuclear magnetic resonance (NMR). Any 
nucleus which possesses a magnetic moment will attempt to align itself 
with the direction of a magnetic field, the quantum alignment being 
dependent, among other things, upon the strength of the magnetic field and 
the magnetic moment. In MRI, a uniform magnetic field B.sub.0 is applied 
to an object to be imaged; hence creating a net alignment of the object's 
nuclei possessing magnetic moments. If the static field B.sub.0 is 
designated as aligned with the z axis of a Cartesian coordinate system, 
the origin of which is approximately centered within the imaged object, 
the nuclei which posses magnetic moments precess about the z-axis at their 
Larmor frequencies according to their gyromagnetic ratio and the strength 
of the magnetic field. 
Water, because of its relative abundance in biological tissues and its 
relatively strong net magnetic moment M.sub.z created when placed within a 
strong magnetic field, is of principle concern in MR imaging. Subjecting 
human tissues to a uniform magnetic field will create such a net magnetic 
moment from the typically random order of nuclear precession about the 
z-axis. In a MR imaging sequence, a radio frequency (RF) excitation 
signal, centered at the Larmor frequency, irradiates the tissue with a 
vector polarization which is orthogonal to the polarization of B.sub.0. 
Continuing our Cartesian coordinate example, the static field is labeled 
B.sub.a while the perpendicular excitation field B.sub.1 is labeled 
B.sub.xy. B.sub.xy is of sufficient amplitude and duration in time, or of 
sufficient power to nutate (or tip) the net magnetic moment into the 
transverse (x-y) plane giving rise to M.sub.xy. This transverse magnetic 
moment begins to collapse and re-align with the static magnetic field 
immediately after termination of the excitation field B.sub.1. Energy 
gained during the excitation cycle is lost by the nuclei as they re-align 
themselves with B.sub.0 during the collapse of the rotating transverse 
magnetic moment M.sub.xy. 
The energy is propagated as an electromagnetic wave which induces a 
sinusoidal signal voltage across discontinuities in closed-loop receiving 
coils. This represents the NMR signal which is sensed by the RF coil and 
recorded by the MRI system. A slice image is derived from the 
reconstruction of these spatially-encoded signals using well known digital 
image processing techniques. 
B. Local Coils and Arrays 
The diagnostic quality or resolution of the image is dependent, in part, 
upon the sensitivity and homogeneity of the receiving coil to the weak NMR 
signal. RF coils, described as "local coils" may be described as resonant 
antennas, in part, because of their property of signal sensitivity being 
inversely related to the distance from the source. For this reason, it is 
important to place the coils as close to the anatomical region-of-interest 
(ROI) as possible. 
Whereas "whole body" MRI scanners are sufficiently large to receive and 
image any portion of the entire human body, local coils are smaller and 
therefore electromagnetically couple to less tissue. Coupling to less 
tissue gives rise to coupling to less "noise" or unwanted biologically or 
thermally generated random signals which superimpose upon the desired MR 
signal. The local coils may be of higher quality factor (Q) than the body 
coils due to their smaller size. For all of these reasons, local coils 
typically yield a higher signal-to-noise S/N ratio ratio than that 
obtainable using the larger whole body antenna. The larger antenna is 
commonly used to produce the highly homogenous or uniform excitation field 
throughout the ROI, whereas the local coil is placed near the immediate 
area of interest to receive the NMR signal. The importance of accurate 
positioning leads to the development of local coils which conform to the 
anatomy of interest, yet function to permit ease of use. 
While the smaller local coil's size works to an advantage in obtaining a 
higher S/N ratio, this reduced size also presents a disadvantage for 
imaging deep-seated tissues. Typically, the single- conductor coil 
diameter which yields the optimal S/N ratio at a depth `d` is a coil of 
diameter `d` ; hence, larger diameter single-conductor coils are required 
to image regions in the abdomen and chest of human patients. This 
increased coil size results in less than desirable performance, both in 
terms of S/N ratio and homogeneity of the sensitivity profile (which 
effects the uniform brightness of the image), and offers little advantage 
over the body coil of the system. 
The S/N ratio of the NMR signal may be further increased by orienting two 
coils, or coil pairs about the imaged object so that each detects RF 
energy along one of a pair of mutually perpendicular axes. This technique 
is generally known as quadrature detection and the signals collected are 
termed quadrature signals. 
The outputs of the quadrature coils are combined so as to increase the 
strength of the received signal according to the simple sum of the output 
signals from the coils. The strength of the noise component of these 
signals, however, will increase only according to the square root of the 
sum of the squares of the uncorrelated noise components. As a result, the 
net S/N ratio of the combined quadrature signals increases by 
approximately .sqroot.2 over the S/N ratio of the individual coils. 
The quadrature orientation of the two coils introduces a 90.degree. phase 
difference between the NMR signals detected by these coils. Therefore, 
combining the outputs from the two quadrature coils to achieve the above 
described signal-m-noise ratio improvements requires that one signal be 
shifted to have the same phase as the other signal so that the amplitudes 
of the signals simply add in phase. 
The approximate net gain of .sqroot.2 in S/N ratio is achievable primarily 
due to the lack of inductive coupling between the coil pairs. This ensures 
that only the uncorrelated noise components add, in lieu of both the 
uncorrelated and correlated noise components, to reduce the effective S/N 
ratio. Inductive isolation is achieved by geometrically orienting the coil 
conductors such that the mutual inductance is minimized between the coil 
pairs according to the following: 
##EQU1## 
where M represents the mutual inductance between coils 1 and 2 and the 
vector components dl.sub.1, and dl.sub.2 represent segments of coils 1 and 
2 with current amplitudes I.sub.1 and I.sub.2. The denominator represents 
the magnitude difference of the position vectors of each dl segment. The 
condition wherein M is approximately zero with respect to the individual 
self inductances of coils 1 and 2, is known as inductive isolation between 
the coils. 
A method of increasing the S/N ratio of the NMR signal over a larger region 
is to digitally add the post processed signals derived from more than one 
coil; each sensitive to the precessing nuclei within overlapping volumes. 
If two coils' signals are processed and converted into image data 
separately and then added digitally, one can obtain an increase in S/N 
ratio within the larger volume. Separate amplifiers, analog-to-digital 
converters, and image processor channels represent an alternative 
configuration for processing the two signals in lieu of a single 
quadrature combiner. A system of four channels whose signals are derived 
from an array of four coils is described in U.S. Pat. No. 4,825,162. In 
the '162 patent, an array of coils is described wherein the adjacent coils 
overlap to prevent nearest-neighbor interaction (inductive coupling). The 
interaction between the next-nearest-neighbor is supposedly reduced by 
connection of each coil of the array to low input impedance preamplifiers. 
The problem with this solution is, among other things, the use of 
preamplifiers with low input impedance. This additional circuitry is 
costly and adds another set of possible failure modes into the system. 
This preamplifier circuitry is sensitive to coil impedance changes 
resulting from patient loading variations as well as to noise spikes or 
power surges within the receiver chain. 
One can minimize the effects of next-nearest-neighbor interaction if one 
properly utilizes the formulation above and in the following arguments to 
minimize inductive coupling between all resonant structures. In this case, 
the additional preamplifier circuitry is no longer required. 
First, nearest-neighbor or adjacent coil interaction is a much more 
dominant coupling than the next-nearest-neighbor coupling--usually one or 
two orders of magnitude larger depending upon coil size and spacing. 
Second, if near-neighbor coupling has not been sufficiently minimized, 
then next-nearest neighbor coupling will occur via neighbor-to-neighbor 
interaction as strongly as, or stronger than inductive coupling between 
next-nearest-neighbors. Third, next-nearest neighbor interaction 
(inductive coupling) is further reduced towards zero when the 
next-nearest-neighbor coils are dominantly loaded by coupling to patient 
tissues. Such is the case in mid to high field scanners operating above 20 
MHz. The coil's impedance is also dominated resulting from coupling to 
eddy current loops generated within the patient tissues. This is predicted 
from the mutual impedance formulation 
##EQU2## 
where Z.sub.1d is the driving or output impedance of coil 1, Z.sub.11 is 
the self-impedance of coil 1, (I.sub.2 /I.sub.1) is the ratio of induced 
eddy currents (loop 2) to the current in coil 1, and Z.sub.12 is the 
mutual impedance between the loops which is equal to the radian frequency 
times the mutual inductance between said loops. 
The implication from the above three facts is as follows. If one ensures 
consistent and dominant loading of the coil elements and if one ensures 
that near-neighbor coupling has been minimized (that is, inductive 
isolation has been achieved), and if the antenna element size, geometrical 
orientation, and spacing is designed so as to minimize next-nearest 
neighbor coupling, then the array will work properly with little degrading 
interaction amongst the elements. 
Inductive isolation is achieved by geometrically orienting two coil 
conductors such that their mutual inductance is minimized according to the 
above formulation. The condition wherein M is approximately zero with 
respect to the individual self inductances of coils 1 and 2, is known as 
geometric isolation between the coils. This is a special case of inductive 
isolation but is restrictive in application, as discussed below. 
As the coil geometries are sufficiently large or close to the surrounding 
system conductors (antenna, faraday screen, cryostat tubing, etc.) in 
addition to the biological conducting medium, this coupling formula must 
be extended to include M=M.sub.12 +M.sub.13 +M.sub.23 ; where the first 
term on the right side of the equation is as described above, and the 
latter two terms define the coupling resulting from each coil's coupling 
to eddy current loops (loop 3) generated on or within the surrounding 
conductors (system or biological). These additional coupling terms must be 
accounted for in the adjustment of conductor geometries with respect to 
each other spatially. With these terms taken into account, the proper 
critical spacing may be found between coil loops 1 and 2. 
Phased array coils such as described in U.S. Pat. Nos. 4,825,162 and 
5,198,768 only utilize linear coil technology to create an array of coils; 
each sensitive to one vector component of the NMR signal in a unique 
imaging volume to create a coil set with a large ROS. Both of those 
patents focus on obtaining a larger region of sensitivity using a bank of 
coils whose signals input to separate preamplifiers and digital 
reconstruction ports on a computer. This does not produce the optimal S/N 
ratio from within the imaging volumes such as would be the case where two 
coil elements are sensitive to orthogonal vector components of magnetic 
flux within each volume (i.e. an array of quadrature elements). The prior 
art is reliant upon geometric isolation and/or low impedance preamps with 
no compensation for eddy current-induced coupling. The prior art also 
restricts the geometry to a planar surface only. This restriction is due, 
in part, to the fact that the prior art is dependent upon geometric 
isolation only, and this alone is inadequate to ensure sufficient 
isolation between non-adjacent pairs of non-planar conductors. This type 
of coil array presents engineering challenges in maintaining isolation 
between elements as additional elements are co-located in space, thereby 
increasing the potential for inductive coupling. 
SUMMARY OF THE INVENTION 
It is an object of the present invention to provide an improved NMR local 
coil designed to conform to a patient's particular anatomical region and 
place antenna conductors within close proximity to the entirety of said 
anatomy. 
Another object is to provide an antenna geometry which permits ease of 
positioning on or about the patient anatomy. 
Still another object is to provide a set of coil conductors which closely 
couple their regions of sensitivity to a desired region of a patient in a 
quadrature fashion. 
Another object of the present invention is to provide improved electronic 
configurations of coil conductors which yield a higher signal-to-noise 
ratio over a larger ROS than other configurations. 
Another object is to provide an array of quadrature coil elements which 
will interface with a multiple-channel receiver system and yield an 
improved S/N ratio over a very large region of interest. 
The NMR Quadrature Detection Array of the present invention includes two 
generally linear arrays; one oriented in a horizontal plane and sensitive 
to vertically oriented changing magnetic flux, and the other oriented in a 
vertical plane bisecting the horizontal plane and sensitive to 
horizontally oriented flux changes. Each horizontal coil is electronically 
paired with a vertical coil such that their NMR signals are combined in a 
quadrature fashion before transmission to the NMR system preamplifier. The 
coil arrays are sized to accommodate easy positioning within the NMR 
system table top assembly as well as to facilitate rapid and easy patient 
positioning about the coil elements. The legs of a patient rest on top of 
the horizontal section while straddling the vertical section; hence 
quadrature detection is possible over the entire length of both legs, and 
sections of the coil can be activated or de-activated for imaging smaller 
or larger fields-of-view. 
The foregoing and other objects and advantages of the invention will appear 
from the following description. In the description, reference is made to 
the accompanying drawings which form a part hereof and in which there is 
shown by way of illustration, a preferred embodiment of the invention. 
Such embodiment does not necessarily represent the full scope of the 
invention, however, and reference must be made therefore to the claims 
herein for interpreting the scope of the invention.

DESCRIPTION OF THE PREFERRED EMBODIMENT 
Referring now to the drawings, in which similar or corresponding parts are 
identified with the same reference numeral, and more particularly to FIG. 
1, the NMR quadrature detection lower extremity array of the present 
invention is designated generally at 10 and includes a generally 
vertically oriented segment 12 mechanically connected to a generally 
horizontally oriented posterior segment 14, to form an inverted "T" shape. 
Referring now to FIG. 2, vertical segment 12 includes a plurality of 
electronic NMR coils 16, 18, 20, 22, 24 and 26 enclosed within a durable 
plastic housing 15 (shown in FIG. 1). Each of these coils is resonant to 
the NMR signal frequency and inductively isolated from other adjacent and 
non-adjacent vertical coils by way of adjustment of coil overlap, with 
compensation for surrounding eddy current-induced coupling effects, and by 
appropriate coil sizing to minimize cross-talk of non-adjacent pairs. Note 
in FIG. 3 that the coil loops (loops 16 and 18 for example) have mitered 
corners 17 at 45 degrees such that the overlapping conductors have minimal 
overlapping surface area. Minimizing this area minimizes the capacitive 
coupling which exists between adjacent conductors; hence further improving 
intercoil isolation. The mitered corners also prevent conductor segments 
from adjacent coils, which are in extremely close proximity to one 
another, from having parallel currents. Forcing these proximal segments to 
be orthogonal to one another also minimizes this inductive coupling 
mechanism. 
Similarly, horizontal segment 14 includes a plurality of electronic NMR 
coils 28, 30, 32, 34, 36, and 38 enclosed within a common plastic housing 
15 (shown in FIG. 1) with the coils of the vertical segment 12. Each of 
the horizontally opposed coils is also inductively isolated from other 
adjacent (via overlaps such as 40) and non-adjacent horizontal coils as 
discussed above. Geometric isolation via precise centering of segment 12 
with respect to segment 14 is all that is required to inductively isolate 
the vertical coils of segment 12 from the horizontally opposed coils of 
segment 14. 
Referring now to FIG. 4, each vertical coil (V1-V6) is paired with the 
adjacent horizontal coil (H1-H6) via electronic connection to a quadrature 
combiner circuit (C1-C6). Each of the six coil pairs is identical in as 
far as general electrical schematic so the following discussion will focus 
on one coil pair V 1 and HI, 26 and 38 respectively. Coil 26 develops NMR 
signal across junction capacitors 42 and 44 resulting from changing 
horizontally oriented magnetic flux components generated by the NMR 
system. The impedance across either of the capacitors 42 or 44 is matched 
to the impedance of 50-ohm transmission line 46 via a matching circuit 48. 
Referring now to FIG. 5, the electrical schematic of circuit board 48 is 
shown. Since the component scheme is identical to that of all matching 
boards, designated with M on FIG. 3, only circuit board 48 will be 
described in detail. Series capacitors 42 and 44 are sized to 
appropriately resonate the coil loop 26. This series configuration reduces 
the total impedance across the capacitor 44 which must be impedance 
matched to the 50 ohm transmission line 46. Inductors 70,72,74, and 76 and 
capacitor 78 form a "modified Tee" impedance matching network which 
matches the complex impedance developed across capacitor 44 to the 50 ohm 
transmission line 46. Inductors 70 and 74 are series connected between 
capacitors 42 and 44 and one conductor 80 of transmission line 46. 
Inductors 72 and 76 are series connected between the opposite side of 
capacitor 44 and a second conductor 82 of transmission line 46. One 
terminal of capacitor 78 is connected between inductors 70 and 74, while 
the other terminal of capacitor 78 is connected between inductors 72 and 
76, to form the "modified Tee" . Whereas inductors 70 and 74 in 
conjunction with capacitor 78 would be the "standard Tee" configuration, 
the additional inductors 72 and 76 modify the standard Tee and serve two 
purposes. First, they are designed to be approximately equal in inductance 
to inductors 70 and 74; hence creating a balanced-to-unbalanced impedance 
transformer ("balun" to those versed in radio frequency (RF) electronics). 
Second, dividing the total circuit inductance evenly onto both conductors 
80 and 82 of the transmission line 46 also keeps both nodes 84 and 86 
above earth ground so as to prevent establishment of an RF ground within 
the sensitive receiver system. Such an RF ground can produce undesirable 
effects upon the magnetic field homogeneity within the NMR system. 
Diode 88 is connected at one end between inductors 70 and 74, and at the 
other end to node 86, and serves as a decoupling diode which is activated 
by system-provided direct current (DC)--voltage pulses. During the 
transmit mode of the MRI data acquisition cycle, a DC voltage forward 
biases diode 88 into a conduction state; hence, effectively placing 
inductor 70 in parallel with capacitor 44. Together these components 
create a high impedance circuit to the RF currents induced upon the 
resonant coil structure, loop 26, thereby decoupling the coil loop from 
the transmit antenna power. 
Referring again to FIG. 4, coil 38 develops NMR signal across junction 
capacitors 50 and 46 resulting from changing vertically oriented magnetic 
flux. Similarly, the impedance across capacitor 46 is matched to the 
impedance of transmission line 54 via matching circuit 56. Cables 46 and 
54 carry NMR signal to the combiner circuit 58 (C1) for phase shifting and 
quadrature addition before the combined output is transmitted to the 
system preamplifier via cable 60. 
All coil pairs (V2, H2; V3,H3, etc.) operate identically to that of pair 
V1, H1 (coils 26 and 38), and all matching boards (denoted with M) are 
identical in operation as are all combiner boards (C1-C6). 
All transmission lines, 60-65, from the combiner (C1-C6) outputs are 
identical in length to maintain the same NMR signal phase shift between 
the coils in the array and the system preamplifier. This is to ensure 
proper digital encoding and further processing in the time domain. Note 
that to ensure equal lengths of transmission lines, non-overlapping `S` 
loops 59 are made in the cables connecting the combiners which are located 
closer to the preamplifier. 
Referring now to FIGS. 6 and 7, iso-magnetic flux contour plots 90 and 92 
of both a vertical coil 26 and a horizontal coil 38, respectively, are 
shown superimposed upon the cross-section of human legs 94. Note that the 
iso-flux contours of the figures represent the same unit of magnetic field 
per unit current upon the coil conductor (H/I), and that each contour line 
is plotted with the same incremental change in comparison to the adjacent 
contours. The corresponding direction of the equal-magnitude contours 90 
and 92 at their points of intersection within the human tissues reveal the 
degree of orthogonality of the two coil's (26 & 38) sensitivity profiles. 
At points where there exists equal magnetic flux contours at right angles 
to each other, there is maximum quadrature gain. A computer program has 
been utilized to maximize the number of points where this condition occurs 
within the volume of the human tissue, and within geometrical constraints 
such as coil size and shape limitations. 
FIG. 8 represents the quadrature combination of the two coil's (26 & 38) 
fields within the same geometric region as in FIGS. 6 and 7. Note that we 
have optimal quadrature gain throughout the majority of the phantom 
volume. This condition holds over the entire length of the array of FIGS. 
1 and 2; hence an array of quadrature detection coils is presented. 
Whereas the invention has been shown and described in connection with the 
preferred embodiments thereof, it will be understood that many 
modifications, substitutions and additions may be made which are within 
the intended broad scope of the appended claims. There has therefore been 
shown and described an improved NMR local coil which accomplishes at least 
all of the above stated objects.