Extra slice spin tagging (EST) magnetic resonance imaging for determining perfusion

A magnetic resonance technique to measure perfusion of fluid to tissue by inverting all the spins in fluid water hydrogen nuclei of the subject with a non-spatially-selective radiofrequency (RF) pulse and immediately reinverting the spins in an imaging plane of interest such that the spins in the imaging plane experience minimal perturbation of their magnetization while the spins outside the plane (extraslice) are inverted (tagged). To account for signal loss effects of magnetization transfer (MT), a control image is obtained by applying two non-selective inversion pulses and creating a perfusion-weighted image by subtracting the tagged image from the control image. When this tagged image is subtracted from a "control" image, the resulting intensity arises solely from the tagged spins that have flowed into the imaging plane. This resultant image is referred to as a perfusion weighted image, because its intensity is a function of the rate of perfusion into the imaging plane. The intensity is also a function of the relaxation rate of the inflowing spins, the partition coefficient of the spins, and the relaxation rate of the stationary tissue. Estimates on these values can be made and can be used to calculate a perfusion map, which is an image that is created in which the intensity of each voxel is proportional to the perfusion of spins into the voxel. EST MRI is entirely noninvasive and overcomes some of the problems associated with the other noninvasive MRI techniques.

FIELD OF THE INVENTION
 The present invention related generally to the field of perfusion
 measurement and more particularly to the field of methods of measuring
 perfusion using non-invasive imaging techniques.
 BACKGROUND OF THE INVENTION
 Perfusion of blood to tissue is extremely important to cell and organ
 viability. Lack of blood flow, or ischemia, can lead to the deleterious
 consequences associated with stroke, myocardial infarction, kidney
 failure, pulmonary embolism, avascalar necrosis of the hip, etc. There are
 also abnormal conditions that may result in increased blood flow that
 would be useful to noninvasively assess, such as is found in certain
 tumors, uterine fibroids, avascular malformations, and the like.
 There are a number of methods that have been developed over the years to
 determine amounts of blood flow using freely diffusible. A nonradioactive
 isotope of xenon has been used as a contrast agent in X-ray Computed
 Tomography (CT). A radioactive isotope of xenon has been radionuclide
 tracer studies. However, xenon is an anesthesia and must be used with
 caution. Radioactive isotopes of oxygen and fluorine have been used for
 assessment of blood flow via PET imaging techniques. Although these
 techniques are useful, their invasive nature has limited widespread
 implementation. Magnetic resonance imaging (MRI) methods offer the
 possibility of noninvasively determining perfusion.
 MRI has traditionally used exogenous contrast agents injected intravenously
 in order to measure blood flow. The most commonly used agents are chelates
 of metals (gadolinium and manganese) that enhance longitudinal (T1)
 relaxivity and result in bright areas on T1 weighted images. The change in
 intensity can be used to calculate perfusion rates. The most commonly used
 MRI contrast agent is gadolinium diethylenetriamine pentaacetic acid
 (Gd-DTPA). This is introduced into the bloodstream by intravenous
 injection. As the coast agent flows into the tissue area where an image
 will be acquired (the imaging plane), it produces a unique signal that can
 be imaged through a variety of image acquisition techniques. Using
 contrast agents has a number of drawbacks, however, including: 1) the need
 for rapid imaging to capture the first pass of the contrast agent before
 it enters the veins, which could produce signal that can interfere with
 the arteries; 2) the need for accurate timing to begin the image
 acquisition to ensure bright arterial and dark venous signal; 3) inability
 to signal average to improve image quality due to the rapidly moving
 contrast agent; 4) possible adverse patient reaction to the contrast
 agent; and 5) cost of the contrast agent, especially in high-dose
 perfusion imaging experiments, in which the use of double- or triple-dose
 contrast could add hundreds of dollars to the cost of the exam.
 MRI methods have been developed that are based on the use of water in the
 blood as an endogenous agent to produce image contrast. The advantage of
 this class of methods is that these techniques are entirely noninvasive
 and repeated measurements can be made for long periods of time. One of the
 first attempts to use blood water as an endogenous tracer was by Denis Le
 Bihan, et al. who used large magnetic field gradients to sensitize the
 image to small motions, including diffusion and perfusion. The problems
 associated with Le Bihan's method is that it is extremely sensitive to
 bulk motion and that it is difficult to segregate the contributions of
 perfusion to changes in signal intensity.
 Another class of techniques involves measuring the signal difference
 between a image acquired after excitation ("tagging") of spins outside an
 imaging plane and an image acquired without exciting the spins outside the
 imaging plane. The signal difference between these two images is due to
 perfusion of tagged spins from outside the imaging plane to inside the
 imaging plane. This class of techniques is commonly referred to as
 Arterial Spin Tagging (AST) or Arterial Spin Labeling (ASL).
 This class of imaging techniques creates contrast through the use of
 strategies that tag spins that subsequently flow into the imaging plane.
 The resulting image data can be used to produce either a
 perfusion-weighted image or can be used to calculate a perfusion map where
 intensities are proportional to flow in ml/100 g.sub.tissue min. The
 mathematics used to produce the perfusion maps is derived from
 steady-state equations for the kinetics of a freely diffusible tracer.
 One of the original AST methods is described in U.S. Pat. No. 5,402,785 of
 Leigh, et al. and in the articles of Detre, et al. and Williams, et al.
 This imaging sequence tags spins flowing through a plane proximal to the
 imaging plane either with an inversion (180.degree.) or saturation
 (90.degree.) RF pulse. The spins flow into the imaging plane and decrease
 the signal. The effect of perfused spins into the imaging plane is
 determined by subtracting the tagged image from a control image. The
 tagged image not only has signal decrease from tagged blood, but from
 magnetization transfer (MT) saturation caused in the imaging plane by the
 off-resonance tagging pulse. In order to account for this, a "tagging"
 pulse is applied an equal distance on the distal side of the imaging
 plane.
 There are several drawbacks to this method. (1) The RF pulses must be
 applied for long period in order to allow the tagged spins to reach a
 steady-state in the imaging plane. This can lead to SAR (specific
 absorption rate) problems. (2) As the spins are being tagged in the distal
 plane, previously excited spins are losing their tag in the imaging plane
 due to longitudinal T1 relaxation. (3) The tagging plane must be carefully
 placed in order to tag inflowing spins. This is not so hard to do in the
 head where tagging can be performed using a axial plane through the
 carotid arteries. Tagging the heart and other parts of the body would be
 difficult. (4) The tagging in the control image can result in the
 excitation of spins in the venous blood.
 U.S. Pat. No. 5,846,197 of Edelman, et al., and the article of Edelman, et
 al., describe another method called Signal Targeting with Alternating
 Radiofrequency (STAR) which is similar to Leigh's method but uses a
 non-steady state tagging of arterial spins by inverting a thick slab
 proximal to the imaging plane proximal to the imaging plane. The tagged
 image is acquired and subtracted from a control image in which no tagging
 is performed. The disadvantages of this technique are similar to Leigh's
 method in that the tagging plane must be carefully placed. Another
 disadvantage of STAR is that it did not (as originally proposed) properly
 account for signal loss due to MT effects.
 The other category of spin-tagging techniques involve tagging spins on a
 "slice" of tissue centered on the imaging plane. Existing methods are FAIR
 (Flow-sensitive Alternating Inversion Recovery) by Kim et al.; UNFAIR
 (Uninverted Flow-sensitive Alternating Inversion Recovery, by Helpern et
 al.; and FIBB (Functional Imaging with pulsed Black Blood).
 FAIR (S. G, Kim) is an imaging method that acquires two images: One
 following a spatially selective RF pulse and one following a nonselective
 pulse. The difference between the two images yields a perfusion weighted
 image. One of the main drawbacks with FAIR is the dependence on the order
 of image subtraction between the control image and the tagged image. FAIR
 has all spins in the imaging plane inverted. The tissue spins in the
 imaging plane relax back towards a ground state. In the tagged image,
 water spins that are fully relaxed from outside the imaging plane flow
 into the tissue and cause the tissue to relax faster than in the control
 image (in which the water spins from outside the imaging plan are
 inverted). Different issues have different T.sub.I, which is to say they
 will relax at different rates. After the initial inversion pulse, spins
 are inverted and have a large, negative signal. As they relax, they pass
 through the x-y plane, at which point they produce no signal. They
 continue to relax until they align with the positive z-axis, and have
 maximal positive signal. Thus the signal goes big to small to big.
 Different tissues do this at different rates. One advantage that EST has
 over FAIR is that the order of image subtraction does not change as a
 function of time after inversion (Ti), i.e., in EST the tag image is
 always subtracted from the control image. In areas where there are tissues
 with different longitudinal relation times (T.sub.I), (such as the brain
 with gray and white matter), there are values of T.sub.I where one cannot
 choose a proper order of image subtraction for FAIR; One order will
 improperly zero the gray matter and the other order will improperly null
 white matter. EST does not have this problem.
 Another method, UNFAIR (Helpern et al.) as originally proposed uses two
 selective inversion pulses, one of which is wider than the other. This
 leaves the stationary tissue at equilibrium values and the spins of the
 water inverted in the tagged image. Use of a wide selective tagging pulse
 introduces potential error into image data, particularly for fast-flow
 tissue such as kidney. Perfusion is observed by observing spins from the
 periphery of the imaging plane (inverted by a wide selective pulse) that
 have flowed into the imaging plane (whose spins are reinverted by the
 narrow selective pulse). Error will occur when equilibrium spins of fluid
 outside the wide, inverted area flow into the imaging plane. Such untagged
 equilibrium spins will have a signature nearly identical to the reinverted
 spins in the imaging plane, this will cancel out some of the signal of
 inverted spins flowing into the plane. Thus, UNFAIR may show less
 perfusion than actually exists. EST avoids this problem by inverting all
 spins outside the imaging plane. Another disadvantage of the UNFAIR method
 is that it acquires one line of k-spacer per tag, which limits the amount
 of data and thereby the spatial resolution that can be obtained in a
 reasonable amount of time. This method has only been demonstrated to using
 a 64.times.64 matrix. This 64.times.64 matrix required 64 separate
 excitations, followed by time to allow the spins to relax (referred to as
 the TR). TR is typically 3-4 seconds, so this low resolution image would
 require 3-4 minutes. A 256.times.256 image with 4 signal averages would
 require close to 70 minutes per slice, which is too long to be clinically
 useful.
 REFERENCES
 1. Detre J A, Leigh J S, Williams D S, Koretsky A P, Perfusion Imaging,
 Magn Reson Med, 23(1):37-45, 1992.
 2. Williams D S, Detre J A, Leigh J S, Koretsky A P. Magnetic resonance
 imaging of perfusion using spin inversion of arterial water. Proc Natl
 Acad Sci, 89:212-216, 1992.
 3. Edelman R R, Siewert B, Darby D G, Thangaraj V, Nobre A C, Mesulam M M,
 Warach S. Qualitative mapping of cerebral blood flow and functional
 localization with echo-planar MR imaging and signal targeting with
 alternating radiofrequency. Radiology, 192:513-520, 1994.
 4. Kim S-G, Quantification of relative cerebral blood flow change by
 flow-sensitive alternating inversion recovery technique; Application to
 functional mapping. MRM, 34:293-301, 1995.
 5. D Le Bihan, et al, MRI of intravoxel incoherent motions. Applications to
 diffusion and perfusion in neurological disorders. Radiology, 161:401-7,
 1986.
 6. Helpern J A. Branch C A. Yongbi M N. Huang N C. MR perfusion imaging in
 human brain using the UNFAIR technique. Un-inverted flow-sensitive
 alternating inversion recovery. Magnetic Resonance Imaging 15(2):135-9,
 1997.
 BRIEF SUMMARY OF THE INVENTION
 This invention is called Extraslice Spin Tagging (EST) MRI because it uses
 a series of spatially selective pulses to "tag" spins outside an imaging
 plane. These spins flow into the imaging plane over a few second period
 and decrease the image intensity. When this tagged image is subtracted
 from a "control" image, the resulting intensity arises solely from the
 tagged spins that have flowed into the imaging plane. This resultant image
 is referred to as a perfusion weighted image, because its intensity is a
 function of the rate of perfusion into the imaging plane. The intensity is
 also a function of the relaxation rate of the inflowing spins, the
 partition coefficient of the spins, and the relaxation rate of the
 stationary tissue. Estimates on these values can be made as described in
 the Detailed Description of the Invention and can be used to calculate a
 perfusion map, which is an image that is created in which the intensity of
 each voxel is proportional to the perfusion of spins into the voxel. EST
 MRI is entirely noninvasive and overcomes some of the problems associated
 with the other noninvasive MRI techniques mentioned above.

DETAILED DESCRIPTION OF THE INVENTION
 This invention describes a method to measure perfusion designated
 Extraslice Spin Tagging (EST) magnetic resonance imaging (MRI). This
 method will be useful in determining perfusion in any substance having a
 well defined fluid supply, such as any tissue having a well defined
 arterial flow (including, but not limited to, the liver, heart, brain,
 muscles, cancerous tissue, and kidney). All questions addressing the scope
 of the present invention may be resolved by reference to the specification
 and the attached claims.
 The intensity in images obtained using magnetic resonance imaging is a
 function of many factors, including nuclear spin density, relaxation rates
 (commonly T1 and T2), motion, and others. The way in which the data is
 collected determines the influence of the various intrinsic factors on the
 image intensity. Unlike X-ray imaging that is sensitive only to electron
 density, MRI can afford clinicians a wealth of information due to its
 reliance on so many tissue characteristics. We have invented a means to
 create MR images sensitive to inflow of spins (commonly the hydrogen on
 water in blood) using noninvasive techniques. These flow sensitive images
 are referred to as Perfusion Weighted Images. We also present the means to
 quantitatively calculate the flow rates on a voxel-by-voxel basis to
 create a perfusion map.
 To obtain perfusion weighted images, two separate images are acquired: A
 Tag Image and a Control Image. The perfusion weighted image is obtained by
 subtraction of the Tag Image from the Control Image. The Tag Image is
 acquired in the following manner (see FIGS. 1 and 2). The nuclear spins
 that reside within the radiofrequency (RF) imaging coil are tagged by the
 application of a non-spatially selective 180.degree. inversion RF pulse
 (see FIG. 1, point A, and FIG. 2A). As soon as possible after this
 (typically 1 ms), a spatially selective 180.degree. re-inversion RF pulse
 is applied to the nuclei in the imaging plane in conjunction with a
 magnetic field gradient (see FIG. 1, point B, and FIG. 2B). As a result,
 the nuclear spins in the imaging plane are returned to their original,
 equilibrium magnetization state. A time delay ensues in order to allow
 time for the tagged spins to flow into the imaging plane (see FIG. 1, time
 period labeled TI). This inversion time (TI) after the tagging period is
 followed by a rapid image collection period. This strategy is important to
 making the technique useable clinically. After the data acquisition, a
 long time period must be used to ensure all spins in the RF coil have
 returned to equilibrium (typically several seconds). If multiple lines of
 data are not collected following each tag, the imaging time becomes
 prohibitively long. We have employed several fast imaging schemes to
 collected data. These include, but are not limited to fast spin echo,
 rapid gradient echo, echo planar, HASTE (half-Fourier acquisition
 single-shot turbo spin echo), echo planar imaging, and GRASE (gradient and
 spin echo).
 As the tagged nuclear spins flow into the imaging plane, the re-inverted
 nuclear spins in the imaging plane are affected in two ways. First, the
 image intensity decreases as a function of the perfusion rate, the
 partition coefficient and the T1 relaxation time. Second, there is a
 slight decrease in the signal due to the on-resonance magnetization
 transfer (MT) between tissue water and macromolecules.
 The Control Image (FIGS. 1 and 3) is acquired using the same procedure
 described above, with the exception that the magnetic field gradient is
 removed during the second 180.degree. inversion pulse at time B (FIG. 1,
 point B and FIG. 3B). As a result, all nuclear spins in the RF coil are
 returned to their original spin states. However, the spins in the imaging
 plane experience the same MT effects as they do during the acquisition of
 the tagged image. As a result, the subtracted image intensity is due only
 to the inflowing spins. Imaging is performed in the same manner as for the
 tagged image (see FIG. 1, point C and FIG. 3C).
 The subtraction of the tag image from the control image produces a
 perfusion weighted image. However, quantitative perfusion values can be
 calculated if more measurements are made. The values are then calculated
 on a pixel-by-pixel basis and result in a perfusion map, which is an image
 where the intensity is proportional to the flow rates. The calculation of
 the perfusion map uses the tag and control images to determine the pixel
 intensities M.sub.b (t) and M.sub.b.sup.o, respectively. A T1 map is then
 determined by running a series of Inversion Recovery centric reordered
 turbo Fast-Low Angle Shot (FLASH, a rapid gradient echo imaging sequence)
 images with increasing TI (10,20,50,100,200,500,1000,2000,5000, and 10000
 ms). Curve fitting is performed on the set of intensity values for each
 pixel to determining a T1 map. A monoexponential curve was assumed.
 Because we used magnitude images, it is not clear whether the minimum
 value in this fitted curve is negative or positive. Therefor the minimum
 was discarded in the fit. But we still have 9 values for the fit, which is
 more than sufficient. Points at shorter TIs than the minimum were negated.
 In order to calculate the perfusion values for each pixel, the Bloch
 equation was solved using an added term for perfusion (Williams, et al.):
 ##EQU1##
 For non-steady-state, we can replace equation (2) into the exponential term
 in equation (1), and it is now:
 ##EQU2##
 Solving for f in equation (3) after some manipulation to replace T.sub.Iobs
 with T.sub.I, we get the equation to calculate perfusion for
 non-steady-state cases:
 ##EQU3##
 Results from this analysis are presented below. The pixel intensifies in
 the perfusion image are equal to the perfusion rates in ml/(100 gmin).
 There are some areas of high intensity from large vessels that we
 threshold out of the final perfusion nap in order to allow for better
 control of window and center of the image.
 EXAMPLE 1
 Extraslice spin tagging was used to determine perfusion in the human brain.
 Perfusion images were obtained for five healthy volunteers (four male and
 one female) with an average age of 30 years. Images were obtained using a
 Siemens VISION 1.5T MR system using a quadrature head coil. The pulse
 sequence used is that described in the Detailed Description of the
 Invention and FIG. 1. Briefly, to obtain tagged images, a non-spatially
 selective hyperbolic secant 180.degree. RF pulse (inversion pulse) was
 applied to the subjects (FIG. 1, point A). The subjects were then
 subjected to a magnetic field gradient, followed by a spatially selective
 18.degree. RF pulse (FIG. 1, point B). The re-inversion pulse was applied
 1 ms after the inversion pulse to allow for magnetic field gradient rise
 time. Subsequently, data was obtained using Rapid Gradient Echo sequence
 (in this case, FLASH sequence) (FIG. 1, point C). Control images were
 obtained in the same manner except that the magnetic field gradient was
 not applied (FIG. 1, point B). The magnetic field gradient applied during
 the tagging pulse is adjusted so that it is two-times wider than the
 imaging plane. This ensures that all the nuclei in the imaging plane have
 their spins re-inverted. To test for complete re-inversion of spins,
 tagged and control images were obtained from stationary phantoms (FIG. 5).
 The tagging pulses were 10.24 ms adiabatic hyperbolic secants. A time
 delay of TI is applied after the tagging pulse to allow the inverted spins
 to flow into the imaging plane.
 Several rapid acquisition methods have been used with EST, including rapid
 gradient echo (RAGE), fast spin echo (FSE) and echo-planar imaging (EPI).
 The results in the current example were obtained using a centrically
 recorded, two dimensional RAGE sequence (called FLASH) with the following
 parameters: TR/TE=11.0/4.2 ms, slice thickness=5 mm, FOV=300 mm,
 matrix=128.times.128, TI=100, 300, 500, 830, 1000, 1500, 2000 and 3000 ms.
 Five measurements were acquired and averaged for the tagged and the
 control images, with a time delay between measurements of 4 seconds.
 T.sub.I maps were acquired with images centrically reordered
 two-dimensional FLASH sequence with TR/TE=11/4.2 ms. An inversion pulse
 was applied, delayed by a time variable TI and then followed by image
 acquisition. The TI values used were 10, 20, 50, 100, 200, 500, 1000,
 2000, 5000 and 10000 ms. Amplitude images were non-linearly fitted to
 produce the T.sub.I map on a pixel by pixel basis. The TI at which a
 minimum intensity occurred (TI min) was determined for each pixel.
 Intensities for TI&lt;.tau. min were negated since the sign of the intensity
 for TI=TI min could not be determined. Pixels in which the non-linear
 algorithm failed would be assigned a 0 value. Pixels below a threshold
 value of 10% of the maximum were not fitted. Images for the T.sub.I
 calculation were performed with NEX=1. Total imaging time for the
 perfusion and T.sub.I data was roughly five minutes per imaging plane.
 The signal differences from the tagged and control images were analyzed to
 determine the perfusion rate according to the equation:
EQU M.sub.c -M.sub.t =[2M.sub.c fTI exp(-TI/T.sub.I)]/.lambda. (5)
 where M.sub.c is the magnitude of the control image (FIG. 6A), M.sub.t is
 the magnitude of the tagged image; f is the perfusion rate is ml 100
 g.sup.-1 min.sup.-1 ; TI is the time delay between the first non-spatially
 selective RF pulse and the spatially selective RF pulse; T.sub.I is the
 longitudinal relaxation rate of the pixel as determined from the T.sub.I
 map; and .lambda. is the blood/tissue partition coefficient. A perfusion
 map was created from this information by using the perfusion value in ml
 100 g.sup.-1 min.sup.-1 for the gray level value in the calculated image
 and is shown in FIG. 6B. This technique allows for the easy assessment of
 flow rates in different parts of the brain using the standard region of
 interest tools on the imaging system.
 All subjects tested demonstrated rapid perfusion to the brain. The grey
 matter perfusion was roughly twice the rate of the white matter. The
 signal from the tagged spins lasted only 3-4 seconds after tagging
 (roughly 3-4 times T1). The average values for perfusion to different
 areas of the brain in ml 100 g.sup.-1 min.sup.-1 +/- standard deviation
 are as follows: global 71.4+/-7.6, gray matter 105.6+/-7.4, and white
 matter 44.0+/-27.8. Complete results for each of the volunteers tested is
 given in Table 1.
 Table 1 represent perfusion values for the five subjects examined. The
 perfusion values are given in ml 100 g.sup.-1 min.sup.-1 ;