Modular whole-body gradient coil comprising first and second gradient coils having linear gradients in the same direction

An MRI system incorporating modular whole body gradient coils that can selectively be used for conventional imaging or for ultra-fast imaging. A central modular coil is used alone for the ultra-fast imaging or used in combination with a modular flanking coil for conventional imaging.

FIELD OF THE INVENTION 
This invention relates in general to magnetic resonance imaging (MRI) 
systems. It more particularly relates to an improved MRI system 
incorporating whole-body gradient coils that can be used either for 
conventional imaging or for ultra-fast imaging. 
BACKGROUND OF THE INVENTION 
Among the basic components of whole-body (MRI) systems are the gradient 
coils. These coils are usually wound on a cylindrical former which is then 
positioned within the bore of the whole-body magnet. The inner diameter of 
the coil former is small enough to enable effective use of the whole-body 
RF coil placed thereon and large enough to enable access to the patient 
who is in the interior of the former. 
As is well known, to perform MRI in both two dimensions and three 
dimensions it is necessary that the coil former contains gradient coils 
appropriate to three orthogonal normal axes, i.e. the X, Y and Z axes. The 
X and Y gradient coils are usually longitudinally symmetrical about the X 
axis and the Y axis respectively. These coils are also separated 
symmetrically by the Y and X axes respectively. That is the X gradient 
coils are bisected by projections of the X axis and the Y gradient coils 
are bisected by projections of the Y axis. The X gradient coils are 
separated and equidistantly spaced from projections of the Y axis. They 
are also equidistantly spaced from the junction of the X, Y and Z axes. 
Similarly, the Y gradient coils are separated and equidistantly spaced 
from the X axis and equidistantly spaced from the junction of the X, Y and 
Z axes. The X gradient coil and the Y gradient coil can be simple coils 
or, more often distributed saddle type coils. The Z gradient coils, on the 
other hand, are separated from and equidistant to the X and the Y axes and 
are symmetrical about the Z axis. The Z gradient coils can be either 
simple pairs or distributed Maxwell pairs of coils. These are the primary 
gradient coils. 
In addition to the primary gradient coils a further set of secondary coils 
wound on a cylindrical former concentric but external to the primary 
coils, is often used in order to reduce the generation of eddy currents 
within the MRI magnetic system. The eddy currents, as is well known, 
degrade image quality. The second set of outer coils are generally 
referred to as "screen coils" or "a screen". 
Theoretical techniques exist to optimize the shape of the flux distribution 
of each set of gradient coils in order to optimize both the linearity of 
the gradient fields over a large volume and the efficiency of the gradient 
coils. In practice, there is always a trade-off between linearity and 
efficiency because of the gradient driver capabilities. Efficiency 
generally means, the energy efficiency; that is the number of gradient 
lines per unit current. 
For conventional imaging it is usual to specify linear gradient fields over 
a central volume that is approximately 50 cm in diameter and has a length 
of up to approximately 60 cm. The linearity of the gradient fields within 
the central volume for conventional MRI are specified to be within plus or 
minus 5%. A large volume in the axial direction is required for certain 
MRI applications such as the imaging of the spine, where linear gradients 
are required over the length of the spine. It is generally sufficient to 
drive such gradient coils with conventional amplifiers that can provide 
250 A at 150 V. This is sufficient since in most conventional scans, 
maximum gradient fields of only 10 mT/m are required with slow rated of 10 
mT/m/ms. 
Suppliers and users of magnetic resonance imaging systems are opting more 
and more for ultra-fast MRI. In ultra-fast MRI systems a complete MRI 
image can be acquired in less than 1 second. A particular technique which 
is used to acquire images in less than 100 ms is the Echo-Planar Imaging 
(EPI) technique which was first suggested by Peter Mansfield, (see his 
U.S. Pat. No. 4,165,479). For such techniques, the performance of the 
gradient coils in providing peak gradient fields and in providing the slow 
rate has to be better--in some cases by a whole order of magnitude--than 
the gradient fields used for regular imaging. A typical requirement for 
EPI gradient performance is 30 mT/m peak gradient with a slow rate of 180 
mT/m/ms. 
When a single gradient coil is used for both EPI and for conventional scans 
serious problems arise. Among other things, the larger the linear volume 
of the coil, the less efficient is the coil, since coil efficiency is 
inversely related to the volume of the coil. In addition, the inductance 
of such coils tend to be large and ultimately limit the slow rate 
attainable with conventional amplifiers. Therefore, when ultra-fast MRI is 
used, the inductance of the large coils tends to degrade the ultra-fast 
operation of the coil and, accordingly, for the ultra-fast operation, 
expensive additional apparatus is required such as special amplifiers and 
special switches, for example. 
One solution to overcoming the loss of efficiency is the use of 
semi-conductor resonant switches. By making the gradient coil part of a 
tuned resonance circuit it is possible to take advantage of the energy 
conserving operation in a controlled manner using semi-conductor switching 
devices and high voltage power supplies. This approach enables the 
requisite performance but on large volume coils in creases the Bpk and the 
dBpk/dt--the peak magnetic field due to the gradient coil and the time 
derivative of the peak magnetic field respectively. For safety purposes it 
is important to restrict the level of Bpk and, therefore, dBpk/dt to which 
the patient is exposed. 
Also, where gradient coils with large linear volumes are used, it is 
unavoidable that the rate of magnetic field change in certain parts of the 
coil is higher than that required for imaging. Accordingly, it is inherent 
in such designs that large components of the gradient field exist outside 
the desired linear region. Depending upon coil design, these large 
components can and generally do exceed the maximum gradient field 
specification for the imaging volume. If Bpk or the derivative with 
respect to time of certain components of these gradient fields reaches a 
high enough value during the scan then it is possible to elicit peripheral 
nerve stimulation in the patient under diagnosis. This is a highly 
undesirable and in fact prohibited side effect. 
It is an object of the present invention overcome the above enumerated and 
other shortcomings that arise when using the same coil for both 
conventional imaging and for ultra-fast imaging in magnetic resonance 
imaging systems. 
Efficiency is inversely proportioned to linear volume so that a way to 
improve efficiency of the gradient coil is to reduce its volume. This can 
be achieved by reducing either the diameter, the length of the linear 
region or both. However, since the patient must fit into the coil it is 
not practical to reduce the diameter. The length of the coil is reducible. 
In addition if the length of the linear volume is limited then the peak 
fields that occur outside the region will also be proportionately lower 
thereby lessening the likelihood of inducing peripheral nerve stimulation. 
A problem, however, with this simple approach is that the facility to 
image anatomy requiring large linear volumes (such as the spine) is 
removed. The gradient coil is then only useful for ultra-fast imaging of 
small regions of interest. 
BRIEF SUMMARY OF THE INVENTION 
To overcome the above enumerated problems in accordance with the present 
invention gradient coils are designed in modular form to provide a 
gradient coil of high efficiency and a tailored usable volume. The modular 
gradient coil comprises a central coil of high efficiency and a limited 
usable volume. To this central coil correction coils are added which 
modify the field profile to increase the usable volume. Thus, the coil 
works in two modes. In a high speed mode just a central portion of the 
coil is enabled to make efficient use of available amplifier power and to 
enable fast switching of gradient fields while limiting the level of 
exposure of the patient to Bpk and rapidly varying differential magnetic 
fields (dBpk/dt's). When the conventional mode of operation is required 
then the central portion of the coil together with correction coils are 
driven to enable large volume coverage for conventional imaging at a 
reduced slow rate. 
More particularly, in accordance with one aspect of the present invention, 
an MRI system is provided comprising: 
a magnet for supplying a homogeneous static magnetic field to align spins 
in a patient positioned in the static magnetic field, 
a radio frequency transmitter for generating signals at Larmor frequencies, 
radio frequency coil apparatus for transmitting said radio frequency 
signals to encompass said patient to tip said aligned spins to have at 
least a projection on a plane normal to the static magnetic field, 
X, Y and Z gradient coils for applying X, Y and Z gradients to the static 
magnetic field to enable position encoding of free induction decay (FID) 
signals that are emitted by the tipped spins after removal of the radio 
frequency signal, 
at least one of said X, Y and Z gradient coils comprising at least one 
modular set of gradient coils, 
said radio frequency coil apparatus also receiving said FID signals, 
signal processing means for processing said FID signals received by the 
radio frequency coil apparatus to obtain image data, 
image display means for displaying images in response to said image data, 
and 
a first modular gradient coil of said at least one set of modular gradient 
coils constructed and arranged to provide a first region having 
substantially linear gradients within said static magnetic field for use 
in ultra-fast MRI, 
a second modular gradient coil of said at least one set of modular gradient 
coils constructed and arranged to provide a second region within the 
static magnetic field having substantially linear gradients for use in 
conventional MRI, said second linear region being larger than said first 
linear region. 
A further feature of the invention couples the first and second of said 
modular gradient coils serially to the same power supply and/or features 
winding both said first and said second of said modular gradient coils on 
the same former. 
Yet a further feature of the invention includes the use of screen coils 
external to said modular gradient coils; said screen coils having a larger 
diameter than said gradient coils and also being modular. 
Still a further feature of the invention includes said first and second 
modular gradient coils having different diameters. For conventional 
operation in imaging, both modular gradient coils are used and for 
ultra-fast operation only the first modular gradient coil is used. When 
both modular gradient coils are used then there is a larger volume wherein 
linear gradients are obtained.

A DETAILED DESCRIPTION OF THE INVENTION 
FIG. 1 at 11 shows a typical MRI system in block diagram form. The main 
magnet 12 is preferably a super-conducting magnet, although within the 
scope of the present invention, the magnet can be of other well known 
types. The magnet 12 contains a bore 13 enabling the entry of a patient 
into the magnetic field. The super-conducting magnet provides a large 
homogeneous static magnetic field. A patient shown at 14 is inserted into 
the bore 13 using bed arrangement 16 so as to be within this large static 
magnetic field. The large static magnetic field causes "spins" in the 
patient to be statistically aligned with the large static magnetic field. 
Encoding signals within the magnet are provided by the gradient coils (not 
shown per se) within the magnet. The X, Y and Z gradient coils are driven 
by X gradient driver 17, Y gradient driver 18 and Z gradient driver 19, 
respectively. The X, Y and Z nomenclature refers to the imaginary 
orthogonal axes shown at 21 used in describing MRI systems; where the Z 
axis is an axis co-axial with the axis of the bore hole. The Y axis is the 
vertical axis extending from the center of the magnetic field and the X 
axis is the corresponding horizontal axis orthogonal to the other axes. 
Radio frequency pulses generated by transmitter 22 and applied through 
multiplexer 23 and radio frequency coil apparatus 24 act to tip the 
aligned spins so as to have a projection, for example, in the X, Z plane; 
the X, Y plane or the Y, Z plane. The spins when realigning after the 
radio frequency pulse is removed generate free induction decay (F.I.D.) 
signals which are received by the radio frequency coil apparatus 24 and 
transmitted through the multiplexer to the receiving circuit 26. From the 
receiving circuit the received signals go through the controller 25 to an 
image processor 27. The image processor works in conjunction with a 
display memory 28 to provide the image displayed on display monitor 29. It 
should be noted that the radio frequency coil apparatus 24 can comprise 
separate coils for transmitting and receiving or the same coil apparatus 
24 could be used for both transmitting and receiving the RF signals. 
FIG. 2a shows the inventive modular gradient coil set 31. Thus the 
illustrated modular gradient coil set 31 is shown as a gradient coil in a 
cross sectional view of a portion of the magnet 12. The modular gradient 
coil set 31 is comprised of modular gradient coils 32 and 34. Screen coils 
38 and 39 are associated with the gradient coils in a manner designed; 
among other things, to reduce eddy currents generated by the gradients. 
In one embodiment the modular gradient coil set includes both a longer 
modular gradient coil 32 extending on the flanks of the set and a 
centrally located smaller gradient coil 34 operating together and used for 
conventional imaging operation. The modular gradient coil 34 provides 
linear gradients in only the central portion of the bore shown as portion 
33 in FIGS. 2a and 2b. The modular gradient coils 32 and 34 together act 
as the modular gradient coil set and provide a comparatively large linear 
region 36. The flanking coil 32 does not have to extend over the entire 
central portion of the magnet since for conventional use the central coil 
34 and the flanking coil 32 co-operate to form the required larger region 
of linear gradients. 
For ultra high-speed imaging, only coil 34 is energized. Coil 32 is 
disconnected. With coil 34 operating at a higher power, i.e. with more 
current going through it, high-speed imaging such as that obtained with 
echo planar imaging sequences is possible even with the power supply 
normally used for conventional imaging. When coil 34 alone is energized 
and used in a ultra high-speed imaging sequence such as an echo planar 
imaging sequence, the current going through coil 34 is much larger than 
for the conventional operation. In addition, the region energized 33, is 
much smaller than the region 36 obtained when both modular gradient coils 
32 and 34 are energized. Fast switching speed is possible because of the 
lower inductance of coil 34 as compared to the inductance of coils 34 
combined with coil 32 or coil 32 alone. 
A feature of the present invention in one embodiment includes the use of 
modular screen coil set 37 on a coil former not shown, which is at a 
larger radial distance than the radial distance used for the operational 
modular gradient coil set 31. The screen coil set is shown at 37 and 
comprises modular coils such as coils 38 and 39 for operation in 
co-operation with the modular gradient coil set 31 comprising coil 32 and 
coil 34 respectively. The screen coils, as is well known, operate to 
restrict eddy currents and stray flux fields. 
FIG. 2b illustrates an arrangement for further improving the efficiency of 
coil 34. As shown its associated modular screen coil 39 is at a greater 
radial distance from the center of the magnet than in the embodiment of 
FIG. 2a (d.sub.2 &gt;di). Thereby the separation between gradient coil 34 and 
screen coil 39 is increased but the overall radius of the complete modular 
coil structure remains unchanged. Thus, the order of the screen coil 
layers can be varied to optimize the efficiency of the coil 34. 
FIG. 2c. The pictorial representation of the set of modular gradient coils 
of FIG. 2c illustrates on exemplary embodiment. Therein the coils 32a and 
32b, which make-up the outer or flanking coil 32, are shown as printed or 
inserted on the outer side of former 35. The inner or central coil 34 is 
represented by portions 34a and 34b printed or inserted on the inner side 
of former 35. An important characteristic is that the length of coil 34 is 
shorter than the length of coil 32. Note that for purposes of clarity, the 
screen coil set 37 is omitted from this drawing. 
FIG. 3a illustrates in solid line form the length of the linear region (1/2 
of the top of the plateau) of the flux generated by central modular coil 
34 when combined with the flanking modular coil 32. The length of the 
linear region of the center modular coil 34 alone is illustrated in dashed 
lines form. The coils 32 and 34 are shown as generating a much larger 
linear region when operating together, whereas coil 34 operating alone is 
shown as generating a much smaller linear region. More particularly, as 
shown in FIG. 3a coil 34 operating alone generates a region that is linear 
only for approximately 15 units, (0.1 meter=10 units) whereas coils 32 and 
34 in combination generate a linear region that extends for approximately 
30 units of length. 
FIG. 3b shows a normalized plot of a maximum magnetic field within the 
patient bore for equal gradient strength of the two modes of operation. In 
particular, it shows that the maximum magnetic field is reduced when 
central coil 34 is used alone (dashed line) compared to coils 34 and 32 
used together (solid line). 
FIG. 4a schematically shows a switching arrangement for selectively 
coupling together the gradient coils to control the size of the region of 
linear gradients. More particularly, a schematic representation of 
switching arrangement 41 for switching the coils is shown. In one position 
of the switching arrangement 41 shown by dashed lines, both the coils 32 
and 34 are operated in series while connected in series to their 
associated screens 39, 38. In the second position of the switches shown by 
solid lines only the coil 34 and its associated screen 39 is energized and 
operated. When only coil 34 is energized the modular gradient coil is set 
for ultra-high speed operation. When conventional operation is desired 
both coils 32 and 34 are used. 
FIG. 4b shows the longer switching arrangement 43 for either connecting 
only the gradient coil 32 for conventional imaging where coil 32 is 
designed to operate alone for conventional imaging; (as shown in FIG. 2b) 
or only connecting coil 34 for ultra-fast operation of the imaging system 
11. 
The gradient coils and the screen coils connected together either in 
parallel or in series within the scope of the invention. 
In operation, for conventional imaging, the modular gradient coil set 31 
includes both gradient coils 32 and 34 and their associated screen coils 
38 and 39. For ultra-fast imaging only gradient coil 34 and its associated 
screen 39 are energized. Within the scope of the invention gradient coil 
32 and its screen coil 38 could be operated alone for conventional 
imaging. In that case coils 32 and 38 would extend to generate more linear 
flux in the central part of the magnet bore. 
In the FIGS. 4a and 4b the screen coils 38, 39 are illustrated as wound in 
a direction opposite to the winding of the operational gradient coils to 
effectively screen the gradient coils. Also, it should be understood that 
the switches are schematically shown and many types of switches including 
computer controlled switches could be used. 
While the invention has been described using an exemplary embodiment, those 
skilled in the art will readily appreciate that many variations and 
modifications may be made in this exemplary embodiment while retaining 
many of the novel features and advantages of this invention. For example, 
while cylindrical coils are illustrated that are planar or elliptical in 
cross section other shapes could be used. Accordingly, all such 
modifications and variations are intended to be included within the scope 
of the appended claims.