.sup.23 NA and .sup.39 K imaging of the heart

Methods for increasing the efficiency of .sup.23 Na and .sup.39 K imaging of biological tissue are provided. For maximum efficiency, the receiver bandwidth is set equal to or less than the quantity N.sub.ro /T*2. Methods for assessing cardiac tissue viability using .sup.23 Na imaging are also provided.

TECHNICAL FIELD OF THE INVENTION 
The field of this invention is magnetic imaging of biological tissue. More 
particularly, this invention relates to .sup.23 Na and .sup.39 K magnetic 
resonance imaging of the heart and the use of such imaging to assess 
cardiac cell viability. 
BACKGROUND OF THE INVENTION 
In many clinical situations, it is critical to determine whether a given 
tissue is viable following an ischemic episode. For example, in cardiology 
the decision to intervene by thrombolytics, percutaneous transluminal 
coronary angioplasty (PTCA), or coronary artery bypass grafting (CABG) is 
made largely on the assumption that the affected myocardium is viable and 
therefore will benefit from the procedure (Bonow, R. O. Identification of 
viable myocardium. Circulation 94:2674-2680, (1996), Hendel, R. C. and 
Bonow, Ro. O. Disparity in coronary perfusion and regional wall motion: 
effect on clinical assessment of viability. Coron. Art Disease 
4(6):512-520, (1993)). Similarly, treatment of stroke patients is strongly 
influenced by available information regarding tissue viability Shimizu, 
T., Naritomi, H., Kuriyama, Y. and Sawada, T. Sequential changes if sodium 
magnetic resonance images after cerebral hemorrhage. Neuroradiology 
34(4):301-304 (1992)). Thus, one of the most important issues regarding 
the management of patients with cardiovascular disease is knowledge of the 
location and extent of injured but viable tissue. 
Extensive clinical experience has demonstrated that one of the best 
approaches for determining viability is to test for normal cell membrane 
function (Bonow, R. O. Identification of viable myocardium. Circulation 
94:2674-2680, (1996), Hendel, R. C. and Bonow, R. O. Disparity in coronary 
perfusion and regional wall motion: effect on clinical assessment of 
viability. Coron. Art Disease 4(6):512-520, (1993)), i.e. to test for 
continued function of the Na.sup.+ -K.sup.+ pump. The most abundant 
natural isotopes of Na and K, .sup.23 Na and .sup.39 K, can be detected by 
magnetic resonance. In principle, it should be possible to use .sup.23 Na 
and .sup.39 K MRI to non-invasively examine cell membrane function and 
therefore viability. Specifically, we have recently shown in an animal 
model that .sup.23 Na image intensity is approximately 100% higher in 
non-viable compared to viable regions following reperfused myocardial 
infarction (Kim, R. J., Lima, J. A. C., Chen, E-L., Reeder, S. B., Klocke, 
F. J., Zerhouni, E. A. and Judd, R. M. Fast 23Na magnetic resonance 
imaging of acute reperfused myocardial infarction:potential to assess 
myocardial viability. Circulation in press:(1997)). Unfortunately, 
however, the in vivo .sup.23 Na and .sup.39 K MR signals are very small. 
The MR sensitivities for .sup.23 Na and .sup.39 K are only 9.2 and 0.051% 
of the .sup.1 H MR sensitivity and that the in vivo concentrations of 
these nuclei are approximately 1,000 times lower than the in vivo water 
proton concentration. The combination of these factors results in .sup.23 
Na and .sup.39 K MR signals which are approximately 22,000 
(1/4.63.times.10.sup.-5) and 2.1 million (1/4.73.times.10.sup.-7) times 
smaller than the standard .sup.1 H signal, respectively. 
Despite the small MR signal, several groups have succeeded in producing in 
vivo .sup.23 Na images of humans Shimizu, T., Naritomi, H., Kuriyama, Y. 
and Sawada, T. Sequential changes if sodium magnetic resonance images 
after cerebral hemorrhage. Neuroradiology 34(4):301-304, (1992))-, Ra, J. 
B., Hilal, S. K., Oh, C. H. and Mun, I. K. In vivo magnetic resonance 
imaging of sodium in the human body. Magn. Reson. Med. 7:11-22, (1988), 
Granot, J. Sodium imaging by gradient reversal. J. Magn. Reson. 
68:575-581, (1986); Granot, J. Sodium imaging of human body organs and 
extremities in vivo. Radiology 167:547-550, (1988), Katz, J. and Cannon, 
P. J. Use of sodium-23 for cardiac magnetic resonance imaging and 
spectroscopy. In: Cardiac imaging, edited by Marcus, M. L., Skorton, D. 
L., Schelbert, H. R. and Wold, G. L. Philadelphia: W. B. Suanders Co., 
1991, p. 828-840; Perman, W. H., Turski, P. A., Houston, L. W., Glover, G. 
H. and Hayes, C. E. Methodology of in vivo humans sodium imaging at 1.5 T. 
Radiology 160:811-820, (1986), Hilal, S. K., Maudsley, A. A., Ra, J. B., 
Simon, H. E., Roschmann, P., Wittekoek, S., Cho, Z. H. and Mun, S. K. In 
vivo NMR imaging of sodium-23 in the human head. J. Comp. Assist. Tomog. 
9(1):1-7, (1985), Winkler, S. S. Sodium-23 magnetic resonance brain 
imaging. Neuroradiology 32:416-420, (1990)). However, most of these groups 
have not attempted to apply recently-developed high speed gradient-echo 
imaging techniques to the .sup.23 Na or .sup.39 K nuclei. One reason for 
the lack of high-speed imaging studies of .sup.23 Na may be that many 
groups are interested in quantifying intracellular Na.sup.+ 
concentrations, for which long TR's to ensure full relaxation are 
desirable. In addition, high-speed imaging of nuclei like .sup.23 Na and 
.sup.39 K is very demanding on gradient hardware due to the low 
gyromagnetic ratios. Recent advances in gradient technology, however, may 
make this issue less significant. 
Because the T.sub.1 and T.sub.2 relaxation times of .sup.23 Na and .sup.39 
K are much shorter than those of .sup.1 H data taken from (Wolf, G. L. 
Contrast agents for cardiac MRI. In: Cardiac imaging, edited by Marcus, M. 
L., Skorton, D. L., Schelbert, H. R. and Wold, G. L. Philadelphia: W. B. 
Saunders Co., 1991, p. 794-810), (Kim, R. J., Lima, J. A. C., Chen, E-L., 
Reeder, S. B., Klocke, F. J., Zerhouni, E. A. and Judd, R. M. Fast 23Na 
magnetic resonance imaging of acute reperfused myocardial 
infarction:potential to assess myocardial viability. Circulation in 
press:(1997), and (Burstein, D., Litt, H. I. and Fossel, E. T. NMR 
characteristics of "visible" intracellular myocardial potassium in 
perfused rat hearts. Magn. Reson. Med. 9:66-78, (1989)) for .sup.1 H, 
.sup.23 Na, and .sup.39 K, respectively!, it is unlikely that direct 
application of fast imaging concepts derived from experience with proton 
imaging would result in optimal imaging parameters for .sup.23 Na and 
.sup.39 K imaging. There continues to be a need in the art, therefore, for 
improved methods for .sup.23 Na and .sup.39 K imaging. 
The present invention provides numerically simulated various imaging 
strategies to maximize .sup.23 Na and .sup.39 K signal acquisition per 
unit time to understand the effect of the short relaxation parameters on 
the data collection. Then using the simulation results as a guide, in vivo 
3D .sup.23 Na images of the human heart were acquired in 15 minutes on a 
modified 1.5 T clinical scanner. The results show that the application of 
high-speed gradient-echo imaging techniques combined with recent advances 
in gradient technology make .sup.23 Na imaging of the human heart 
practical. 
BRIEF SUMMARY OF THE INVENTION 
Numerical simulations of high-speed imaging sequences were developed and 
used to maximize .sup.23 Na and .sup.39 K image SNR per unit time within 
the constraints of existing gradient hardware. The simulation demonstrated 
that decreasing receiver bandwidth at the expense of echo time (TE) 
results in a substantial increase in .sup.23 Na and .sup.39 K image 
SNR/time despite the short T.sub.2 and T.sub.2 * of these nuclei. 
Referenced to the available .sup.1 H signal on existing 1.5 T scanners, 
the simulation suggested that it should be possible to acquire 3D .sup.23 
Na images of the human heart with 7.times.7.times.7 mm resolution and 
.sup.39 K images with 26.times.26.times.26 mm resolution in 30 min. 
Experimentally in humans at 1.5 T, 3D .sup.23 Na images of the heart were 
acquired in 15 min with 6.times.6.times.12 mm resolution and SNRs of 11 
and 7 in the left ventricular cavity and myocardium, respectively, which 
is very similar to the predicted result. The results demonstrate that by 
choosing imaging pulse sequence parameters which fully exploit the short 
relaxation times of .sup.23 Na and .sup.39 K, potassium MRI is improved 
but remains impractical whereas sodium MRI improves to the point where 
.sup.23 Na imaging of the human heart may be clinically feasible on 
existing 1.5 T scanners. 
In accordance with such simulations, the present invention provides a 
method of increasing the efficiency of 23Na magnetic resonance imaging of 
biological tissue, the process including the step of acquiring images with 
a receiver bandwidth equal to or less than R(N.sub.ro /T*.sub.2Na) where R 
is an integer from 1 to about 5. Preferably, R is from about 1 to about 3, 
and more preferably R is from about 1 to about 2. 
In one embodiment, the bandwidth is from about 500 Hz to about 10,000 Hz. 
More preferably, the bandwidth is from about 1,000 Hz to about 5,000 Hz, 
from about 2,500 Hz to about 5,000 Hz and, most preferably, from about 
2,500 Hz to about 3,000 Hz. 
The 23Na images are acquired using a 1.5 T, 3.0 T or a 4.0 T scanner. The 
images are preferably acquired in from about 15 to about 60 minutes, from 
about 30 to about 45 minutes and, most preferably, in about 30 minutes. 
The biological tissue can be located in a living organism. Preferably the 
biological tissue is a human heart. 
In another aspect, a process of this invention increases the efficiency of 
39K magnetic resonance imaging of biological tissue. The process includes 
the step of acquiring images with a receiver bandwidth equal or less and 
R(N.sub.ro /T*.sub.2K), where R is an integer from 1 to about 5. The 
bandwidth is from about 1,000 Hz to about 30,000 Hz and, preferably, is 
about 8,000 Hz. 
The present invention further provides a process of identifying regional 
areas of myocardial damage in vivo. The process includes the step of 
imaging the heart in vivo using .sup.23 Na magnetic resonance imaging 
wherein regions of relative high image .sup.23 Na intensity indicate the 
damaged regions. The damaged regions of the myocardium are nonviable 
tissue resulting from any pathological condition such as ischemia. In 
addition, recently developed rapid gradient-echo techniques used for 
proton imaging have been applied to the sodium nucleus to explore methods 
to reduce .sup.23 Na imaging time to a level which would allow .sup.23 Na 
imaging to become a practical experimental and clinical tool. 
Where the heart is imaged using .sup.23 Na imaging, that imaging is 
accomplished using a) an imaging time of from about 3 minutes to about 20 
minutes; b) about 16 phase encodes per cardiac cycle; c) an echo time of 
from about 2.5 milliseconds to about 6.5 milliseconds; d) a repetition 
time of from about 8 milliseconds to about 20 milliseconds; e) a signal 
averaging of from about 2090 to about 300; f) a voxel size of about 
1.0-1.5.times.2.0-3.0.times.4.0-8.0 millimeters. 
Preferably, the imaging time is from about 6 to about 15 minutes, more 
preferably from about 8 to about 12 minutes and, most preferably about 10 
or 11 minutes. The echo time is preferably from about 3.5 milliseconds to 
about 5.5 milliseconds and, more preferably about 4.5 milliseconds. The 
repetition time is preferably from about 10 milliseconds to about 15 
milliseconds and, more preferably about 13 milliseconds. The signal 
averaging is preferably from about 225 to about 275 and, more preferably 
about 256. A preferred voxel size is about 1.25.times.2.5.times.6 
millimeters. 
The present disclosure establishes the relationship of regional changes in 
sodium image intensity to myocardial viability without the need to 
distinguish between intra- and extracellular Na.sup.+. Regional image 
intensity, both in isolated hearts and in vivo, correlated with myocardial 
viability determined by TTC (triphenyltetrazolium chloride) staining 
techniques, regional differences in sodium content measured using, .sup.23 
Na MR spectroscopy, and regional differences in sodium T.sub.1 and T.sub.2 
relaxation times. 
The physiologic and technical feasibility of using .sup.23 Na MRI to 
examine myocardial viability has been established. Eighteen rabbits 
underwent in situ coronary artery occlusion and reperfusion. .sup.23 Na 
images were obtains in these rabbits, normal rabbits, and normal dogs. In 
infarcted, reperfused regions .sup.23 Na image intensity was greater than 
in viable regions (isolated hearts: 42.+-.5%, p&lt;0.02, in vivo: 95.+-.6%, 
p&lt;0.001). Data from .sup.23 Na MR spectroscopy showed a similar pattern 
(p&lt;0.001). Following acute infarction with reperfusion, a regional 
increase in .sup.23 Na MR image intensity is associated with nonviable 
myocardium. Fast gradient-echo imaging techniques can reduce .sup.23 Na 
imaging time to a few minutes.

DETAILED DESCRIPTION OF THE INVENTION 
Taking the sensitivity, abundance and concentration into account, the 
signal deficits of .sup.23 Na and .sup.39 K are about 22,000 and 2.1 
million, respectively, compared to the .sup.1 H MR signal (Table 1, 
below). 
TABLE 1 
______________________________________ 
In vivo magnetic resonance sensitivities of .sup.1 H, .sup.23 Na, and 
.sup.39 K 
Relative In vivo 
Nucleus 
% Abundance 
Sensitivity 
Concentration 
In vivo signal 
______________________________________ 
.sup.1 H 
99.99 1.000000 100,000 1 
.sup.23 Na 
100.00 0.092500 50 4.63E-05 
.sup.39 k 
93.10 0.000508 100 4.73E-07 
______________________________________ 
For .sup.23 Na and .sup.39 K imaging to be practical in humans, therefore, 
it is necessary to increase the signals by these amounts. As a numerical 
example, consider standard .sup.1 H imaging of the human heart. Using 
state-of-the-art hardware, it is possible to acquire a .sup.1 H image with 
2.times.3.times.10 mm resolution in a single heartbeat, a technique 
typically used in conventional cardiac MRI. The design considerations set 
forth herein outline the method needed to acquire a sodium or potassium 
image with nearly identical SNR as the single heartbeat .sup.1 H image. 
For comparison purposes, the unit proton SNR is defined as that of a 
proton image with 2.times.3.times.10 mm resolution acquired in 1 sec with 
a receiver bandwidth of 100,000 Hz. In a typical 1.5 T clinical scanner 
with these parameters, the proton image SNR in the myocardium was 11. 
Two approaches to increasing .sup.23 Na and .sup.39 K signal are to 
increase voxel size and lengthen imaging time Taking .sup.23 Na imaging as 
an example, increasing the size of the voxel from 2.times.3.times.10 mm to 
6.times.6.times.12 mm would increase signal by a factor of 7. In addition, 
by increasing imaging time to thirty minutes an additional factor of 42 
could be realized (.sqroot.(30 min*60 sec/min)). Although combining these 
two approaches would increase .sup.23 Na SNR by a factor of 294 (=7*42), 
this is still far from the necessary 22,000. 
An additional gain in .sup.23 Na SNR can be realized by employing fast 
imaging pulse sequences which exploit the fast T.sub.1 of .sup.23 Na, such 
as high-speed coherent (FISP, or GRASS) and incoherent (FLASH, or SPGR) 
gradient echo (GRE) sequences run with a very short TR. Furthermore, by 
employing short TRs, fast 3D imaging has significant advantages over a 
comparable 2D sequence for at least two reasons. First, a shorter echo 
time can be realized because the RF excitation is not slice-selective. 
Second, because .sup.23 Na (and .sup.39 K) imaging requires significant 
signal averaging, it is more time efficient to encode a third spatial 
dimension (Ra, J. B., Hilal, S. K., Oh, C. H. and Mun, I. K. In vivo 
magnetic resonance imaging of sodium in the human body. Magn. Reson. Med. 
7:11-22, (1988), Wehrli, F. Fast-scan magnetic resonance: principles and 
applications. Magn. Reson. Q. 6:165-236, (1990)). The entire heart can be 
imaged in the same time as a 2D slice for a constant SNR. 
In order to explore which type of sequence yielded the greatest increase in 
.sup.23 Na SNR, the steady state MR signal for each sequence using the 
T.sub.1, T.sub.2, and T.sub.2 * values given in Table 2, below, for .sup.1 
H, .sup.23 Na, and .sup.39 K were plotted. 
TABLE 2 
______________________________________ 
Gyromagnetic Ratios and Relaxation Times for .sup.1 H, .sup.23 Na, and 
.sup.39 K 
Nucleus 
Gyromagnetic Ratio ( ) MHz/T! 
T1 ms! T2 ms! 
T2* ms! 
______________________________________ 
.sup.1 H 
42.5 870 60 25 
.sup.23 Na 
11.2 35 30 25 
.sup.39 K 
1.98 10 8 8 
______________________________________ 
FIG. 1 shows signal for FISP and FLASH sequences using a TR of 15 ms. As 
expected in the case of .sup.1 H imaging (FIG. 1a), the FISP sequence had 
slightly more signal available. For .sup.23 Na imaging (FIG. 1b) there is 
no real difference between the peak magnitude of the FISP or FLASH signal. 
Contrary to proton-based imaging theory, the FLASH sequence is optimal for 
.sup.39 K (FIG. 1c). 
The major advantage of using GRE imaging is that both the .sup.23 Na and 
.sup.39 K curves realize a much higher signal level compared to protons 
(50% vs. 8% of M.sub.0 for .sup.23 Na and 80% vs. 8% of M.sub.0 for 
.sup.39 K compared to .sup.1 H, respectively). Therefore, choosing the 
optimal sequence type for .sup.23 Na imaging increases the signal compared 
to protons by an additional factor of 6. When combined with larger voxels 
and longer imaging time, the .sup.23 Na signal from fast gradient echo 
imaging is now increased by 1,764 (=294*6). 
Finally, choosing the optimal receiver bandwidth for .sup.23 Na (and 
.sup.39 K) imaging should likely yield an additional increase in signal. 
Although decreasing bandwidth generally increases SNR by reducing noise, 
the optimal choice of bandwidth for .sup.23 Na and .sup.39 K is unclear 
because decreasing bandwidth will result in a longer TE, which will 
decrease signal due to the short T.sub.2 and T.sub.2 * of .sup.23 Na and 
.sup.39 K. In addition, the TR will increase and may affect SNR/time. For 
example, consider the plot in FIG. 2 which shows image SNR per unit time 
for a FLASH sequence as a function of TR with TE=0. As TR decreases, 
SNR/time reaches a plateau at or near the T.sub.1 of each nucleus. One can 
calculate that for 6.times.6.times.12 mm voxels and state-of-the-art 
gradients (25 mT/m), one could achieve a TE of 1.3 ms and a TR of 3.1 ms 
for .sup.23 Na imaging. However, from FIG. 2 increasing the TR to almost 
30 ms results in no penalty in the SNR/time (for TE=0). Therefore, 
neglecting T.sub.2 * losses for the moment, decreasing receiver bandwidth 
from 100,000 (Unit Proton SNR) to 2,500 for .sup.23 Na imaging might 
increase .sup.23 Na signal by as much as 6-fold (=.sqroot.(100,000/2,500). 
Consequently, the combined effects of increasing voxel size, increasing 
imaging time, using a fast 3D imaging sequence, and optimizing receiver 
bandwidth will likely increase the .sup.23 Na signal by 10,584 (=1,764*6), 
i.e. increase the .sup.23 Na signal to within a factor of two of the Unit 
Proton SNR. Analogous arguments suggest that the .sup.39 K signal would 
also approximate the Unit Proton SNR but would require larger voxels. 
Choosing the imaging parameters that maximize SNR/time is difficult without 
full numerical simulation because changing a single parameter such as 
bandwidth directly affects other parameters such as TE, TR, and the 
optimal flip angle. The above considerations, however, make it likely that 
.sup.23 Na and .sup.39 K imaging is possible. Therefore, fast imaging 
pulse sequences for .sup.23 Na and .sup.39 K were numerically tested to 
determine the additional SNR/time which might be made available by 
application of high-speed imaging techniques. 
Pulse Sequence Simulations 
To determine the specific 3D GRE pulse sequence parameters which optimized 
.sup.23 Na and .sup.39 K image SNR per unit time, standard 3D FISP and 
FLASH pulse sequences were numerically simulated. The equations used in 
the simulations (Wehrli, F. Fast-scan magnetic resonance: principles and 
applications. Magn. Reson. Q. 6:165-236, (1990), Vlaardingerbroek, M. T. 
and den Boer, J. A. Magnetic Resonance Imaging, Berlin:Springer-Verlag, 
1996) are given below. 
The 3D FISP and FLASH sequences were simulated as follows. First, the 
variables G.sub.max, SR, .gamma., PW, N.sub.ro, N.sub.pe, N.sub.part, and 
N.sub.el were taken from Tables 2 and 3. For a given bandwidth (BW), the 
dwell time (t.sub.dwell) was calculated as: 
##EQU1## 
the readout time (t.sub.ro) was calculated as: 
EQU t.sub.ro =N.sub.ro .multidot.t.sub.dwell 
and the amplitude of the readout gradient (G.sub.ro) was calculated as: 
##EQU2## 
Next, the number of points by which the echo was offset (N.sub.eo) was 
calculated based on the number of points from the beginning of the readout 
window to the echo (echo location, N.sub.el, from Table 3): 
##EQU3## 
The echo offset was also calculated as a percent of the readout window 
(%.sub.eo): 
##EQU4## 
Next, the time required for the readout dephaser (t.sub.rd), phrase 
enclode (t.sub.pe) and slice enclode (t.sub.s) gradients were calculated 
assuming the gradients are run at maximum: 
##EQU5## 
From these times, the actual encoding time, t.sub.c, was taken as the 
maximum of t.sub.rd, t.sub.pe, t.sub.s. Note that gradient duty cycle was 
not considered because our gradient system is capable of 100% duty. The 
echo time (TE) and repetition time (TR) were calculated as: 
##EQU6## 
Next, the TE and TR values were combined with the Ernst angle (from 
Equation 4) and used with either Equation 2 (FLASH) or Equation 3 (FISP) 
to calculate the MR signal per TR (S.sub.FLASH or S.sub.FISP). Finally, 
SNR/time was calculated using Equation 1. This process was repeated for 
bandwidths varying from 500 to 100,000 Hz to determine the bandwidth which 
maximized SNR/time for .sup.23 Na and .sup.39 K imaging. 
Input parameters to the simulation were: gyromagnetic ratio and relaxation 
times for each nuclei, resolution, maximum gradient strength, maximum slew 
rate, and RF pulse width. 
The simulation required an estimate of T.sub.2 *, which depends on many 
factors such as geometry, field, etc. The T.sub.2 * can be estimated for 
.sup.23 Na and .sup.39 K by assuming a T.sub.2 * for .sup.1 H of 25 msec, 
which is typical on standard 1.5 T systems. Based on this assumption, an 
estimated magnetic field inhomogeneity, .DELTA.B, from the equation 
1/T.sub.2 *=1/T.sub.2 +.gamma..DELTA.B, can be calculated. Given the 
estimated .DELTA.B, T.sub.2 * for .sup.23 Na and .sup.39 K were then 
calculated from their .gamma. and T.sub.2 values and the results are given 
in Table 2, above. 
Because the pulse sequence timing is a strong function of gradient 
performance and resolution, specific imaging parameters given in Table 3, 
below, as typical of human heart imaging on a clinical scanner (1.5 T 
Siemens Vision) were used. 
TABLE 3 
______________________________________ 
Pulse sequence parameters used to simulate .sup.23 Na and .sup.39 K 
imaging 
Nucleus .sup.23 Na 
.sup.39 K 
______________________________________ 
Maximum Gradient (G.sub.max) mT/m! 
25 25 
Slew Rate (SR) (mT/m)/ms! 
25 25 
RF Pulse Width (PW) .mu.s! 
500 500 
Number Readout Samples (N.sub.ro) 
64 64 
Echo Location (N.sub.e) Sample Number! 
12 12 
Number Phase Encoding Steps (N.sub.pe) 
32 32 
Number Slice Encoding Steps (N.sub.part) 
32 32 
______________________________________ 
Given the parameters in Tables 2 and 3, the simulation calculated the 
minimum TE and TR for each pulse sequence type within the constraints of 
the gradient hardware. Once TE and TR were known, SNR per unit time was 
calculated according to Equation 1!, below: 
##EQU7## 
where S is the percent of the fully relaxed magnetization (%M.sub.0) 
sampled per RF pulse. For the case where transverse magnetization was 
spoiled (FLASH or SPGR) (Wehrli, F. Fast-scan magnetic resonance: 
principles and applications. Magn. Reson. Q. 6:165-236, (1990), 
Vlaardingerbroek, M. T. and den Boer, J. A. Magnetic Resonance Imaging, 
Berlin:Springer-Verlag, 1996), S was calculated from the Equation 2!, 
below: 
##EQU8## 
For the case where transverse magnetization was refocussed by rewinding 
the phase and slice encode gradients (FISP or GRASS), S was determined by 
Equation 3!, below (Wehrli, F. Fast-scan magnetic resonance: principles 
and applications. Magn. Reson. Q. 6:165-236, (1990), Vlaardingerbroek, M. 
T. and den Boer, J. A. Magnetic Resonance Imaging, Berlin:Springer-Verlag, 
1996): 
##EQU9## 
For the case of spoiled transverse magnetization, the flip angle (a) was 
taken as the Ernst angle according to Equation 4!, below: 
##EQU10## 
For the case of refocussed transverse magnetization, the optimum flip 
angle was determined numerically by maximizing S as defined in Equation 
3!, above. 
Although it is known that noise decreases with decreasing bandwidth, lower 
bandwidth also increases TE which can result in significant loss of signal 
for .sup.23 Na and .sup.39 K, due to their short T.sub.2 and T.sub.2 *. 
The simulation was, therefore, designed to adjust bandwidth until an 
optimum SNR/t was found. However, as bandwidth is decreased at some point, 
the bandwidth per pixel will exceed the line width, 1/T.sub.2 *, resulting 
in blurring between pixels and a loss of spatial resolution. Bandwidth was 
constrained by the Equation 5!, below: 
EQU Bandwidth&gt;N.sub.ro /T.sub.2 * 5! 
where N.sub.ro is the number of readout samples. 
FIG. 3 shows the simulated effect of receiver bandwidth on optimized image 
SNR/time for .sup.23 Na and .sup.39 K using a voxel size of 
6.times.6.times.6 mm for .sup.23 Na and 12.times.12.times.12 mm for 
.sup.39 K. Note that for each nucleus in FIG. 3, TE, TR and flip angle 
vary with bandwidth according to the simulation. From FIG. 3, for both 
nuclei SNR/time reaches a maximum value at bandwidths between 1,000 and 
10,000 Hz. As previously noted, however, image spatial resolution will 
suffer if the bandwidth is decreased to a point where the bandwidth per 
pixel is less than the line width of the FID (i.e. when bandwidth&lt;N.sub.ro 
/T.sub.2 *). For both nuclei, the bandwidth for which SNR/time was 
maximized was limited by this constraint. The exact location of these 
constraints for sodium and potassium are shown in FIG. 3. Specifically, it 
was found that the optimal bandwidths were 2,560 and 8,000 Hz for .sup.23 
Na and .sup.39 K, respectively, corresponding to dwell times of 390 and 
125 .mu.sec. The important finding demonstrated by the results of FIG. 3 
is that for the parameters in Tables 2 and 3, the optimal SNR/time without 
loss of spatial resolution for both nuclei will occur when bandwidth is 
set equal to N.sub.ro /T.sub.2 *. At those bandwidths, the optimal TR's 
were 26 and 10 ms for .sup.23 Na and .sup.39 K, respectively, and the 
optimal TE's were 5.2 and 2.4 ms, respectively. 
It is interesting to note that for both .sup.23 Na and .sup.39 K the 
maximum SNR/time does not occur at the shortest possible TE (i.e. all 
gradients run at maximum) despite the short T.sub.2 and T.sub.2 * of these 
nuclei (Table 2). To understand this result, consider the .sup.23 Na 
results in FIG. 3 for bandwidths of 100,000 and 2,560 Hz. At a bandwidth 
of 100,000 Hz, the simulation calculated that the gradients could achieve 
a TE=1.3 ms and TR=3.1 ms and at 2,560 Hz the simulation calculated a 
TE=5.2 and TR=26. Because both TRs are shorter than .sup.23 Na T, (Table 
2), the flip angle (24.degree. at 100,000 and 61.degree. at 2,560) was 
adjusted by the simulation according to the Ernst equation (Equation 4) to 
maintain a constant SNR/time with regard to T.sub.1 relaxation (FIG. 2). 
Consequently, the only parameters affecting SNR/time in this example are 
T.sub.2 * decay and bandwidth. Due to the longer echo time, the signal 
decreases by a factor of (e.sup.-1.3/25)/(e.sup.-5.2/25)=1.17. However, 
due to the lower bandwidth the noise is reduced by a factor of 
(1/.sqroot.2,560)/(1/.sqroot.100,000)=6.25. Thus an unexpected result is 
that bandwidth is more important than a short echo time for the parameters 
used in this simulation. 
Another interesting result of the simulation is that given the parameters 
of Tables 2 and 3 and assuming everything is run at the optimal bandwidth, 
one can estimate how large the .sup.23 Na and .sup.39 K voxels would need 
to be to obtain an image SNR equal to the Unit Proton SNR. To make this 
comparison, the relative MR sensitivities given in Table 1 were taken into 
account to allow direct comparison to protons and an imaging time of 30 
minutes was assumed for .sup.23 Na and .sup.39 K. The results are shown in 
FIG. 4. For both nuclei, image SNR increases with voxel size not only 
because the voxels are larger but also because larger voxels decrease 
gradient demands which can affect bandwidth, TE, TR, and flip angle (all 
of which are continuously adjusted by the optimization algorithm). From 
FIG. 4, the simulation predicts that at 1.5 T it should be possible to 
acquire a .sup.23 Na image with 7.times.7.times.7 mm spatial resolution 
and a .sup.39 K image with 26.times.26.times.26 mm resolution in 30 min. 
When compared to a reference proton image (Unit Proton SNR), the simulation 
predicted that .sup.23 Na imaging of the human heart at 1.5 T should in 
principle be possible with 7.times.7.times.7 mm resolution in 30 min (FIG. 
4). Experimentally, a similar result (6.times.6.times.12 mm resolution in 
15 min, FIG. 5). Therefore, despite the fact that the simulation did not 
take into account important but difficult to quantify differences between 
proton, sodium, and potassium imaging such as RF coil performance, the 
simulation provides a useful mathematical framework for evaluating 
strategies for .sup.23 Na and .sup.39 K imaging. 
Potassium Imaging 
Unlike .sup.23 Na MRI, .sup.39 K MRI does not appear to be clinically 
feasible even with the use of optimized imaging parameters. The simulation 
results suggested that human .sup.39 K imaging at 1.5 T in 30 min would 
require voxel dimensions of 26.times.26.times.26 mm (FIG. 4) which is too 
large to be clinically useful. In light of the agreement between the 
simulation and experimental results for sodium imaging, therefore, 
potassium imaging was not evaluated experimentally. Interestingly, SPECT 
and PET image resolutions are typically 10-15 mm (detector limit, full 
width at half maximum). Therefore, the results suggest that if an 
additional 8-fold increase in SNR were available, .sup.39 K MRI should be 
possible in 30 min with a spatial resolution similar to PET and SPECT 
(26.times.26.times.26/13.times.13.times.13 mm=8-fold). However, other 
factors such as "invisible" intracellular potassium Fossel, E. T. and 
Hoefeler, H. Observation of intracellular potassium and sodium in the 
heart by NMR: a major fraction of potassium is "invisible". Magn. Reson. 
Med. 3:534-540, (1986), Pike, M. M., Frazer, J. C., Dedric, D. F., 
Ingwall, J. S., Allen, P. D., Springer Jr., C. S. and Smith, T. W. .sup.23 
Na and .sup.39 K nuclear magnetic resonance studies of perfused rat 
hearts. Discrimination of intra- and extracellular ions using a shift 
reagent. Biophys. J. 48:159-173, (1985)) and substantially different 
design criteria for low frequency RF coils Vlaardingerbroek, M. T. and den 
Boer, J. A. Magnetic Resonance Imaging, Berlin:Springer-Verlag, 1996), The 
ARRL Handbook For Radio Amateurs, Newington, Conn.: American Radio Relay 
League, 1996.) (the .sup.39 K frequency is 2.98 MHz at 1.5 T) make it 
difficult to predict the precise amount of additional SNR needed to make 
.sup.39 K MRI clinically practical. 
Choice of Pulse Sequence 
The results suggest that high speed GRE pulse sequences employing partial 
flip angle RF excitation are best suited for .sup.23 Na and .sup.39 K 
imaging. In addition, 3D rather than 2D imaging appears to have 
considerable benefits including full volume coverage in the same imaging 
time, the ability to post-process the data to yield arbitrary slice 
orientation, and a reduction in TE due to the use of a non-selective 
square RF pulse. Furthermore, the ability to use a non-selective RF pulse 
for 3D imaging was helpful in modifying the 1.5 T clinical scanner for 
.sup.23 Na imaging because it was unnecessary to interface to the Siemens 
hardware for pulse shaping. 
Only rectilinear k-space sampling schemes were studied. Other approaches, 
such as 3D projection reconstruction (Boada, F. E., Christensen, J. D., 
Huang-Hellinger, F. R., Reese, T. G. and Thulborn, K. R. Quantitative in 
vivo tissue sodium concentration maps: the effects of biexponential 
relaxation. Magn. Reson. Med. 32:219-223, (1994), Boada, F. E., Gillen, J. 
S., Shen, G. X., Chang, S. Y. and Thulborn, K. R. Fast three dimensional 
sodium imaging. SMR Proceedings 2:1195, (1995)) have been used for .sup.23 
Na imaging. Projection reconstruction may be particularly useful for 
quantification of sodium concentrations because of the short TE which can 
be achieved Boada, F. E., Gillen, J. S., Shen, G. X., Chang, S. Y. and 
Thulborn, K. R. Fast three dimensional sodium imaging. SMR Proceedings 
2:1195, (1995)). 
Spin echo imaging was not investigated for .sup.23 Na or .sup.39 K for 
several reasons. First, the use of a surface coil for RF transmission is 
difficult for spin-echo imaging, due to an inhomogeneous RF field. Second, 
because the specific absorption rate (SAR) of RF energy increases 
quadratically with flip angle, the SAR for spin-echo imaging would be 
roughly 5 times greater than for gradient-echo imaging. Third, T.sub.2 
relaxation times are so short for .sup.23 Na and .sup.39 K that, at least 
in theory (see Table 2), the T.sub.2 * values for these nuclei are nearly 
the same as T.sub.2 for magnetic field inhomogeneities expected on 
clinical systems. Consequently, signal at a given TE will be nearly the 
same for spin-echo compared to gradient-echo. Furthermore, since spin echo 
has a larger TE because of the 180.degree. pulse, the SNR/time will likely 
be lower. 
Optimal Imaging Parameters 
An important finding of the present study is that the optimal receiver 
bandwidth can be directly calculated from T.sub.2 *. As shown in FIG. 3, 
for both .sup.23 Na and .sup.39 K the bandwidth at which SNR/time is 
maximized is limited by the point at which bandwidth per pixel begins to 
exceed the line width (1/T.sub.2 *). This implies that an important first 
step for .sup.23 Na and .sup.39 K imaging is to collect an FID for the 
particular experimental setup; shim if possible; measure the T.sub.2 * 
from the FID decay; and set the imaging bandwidth to N.sub.ro /T.sub.2 *. 
This finding has two other important implications. First, additional 
signal may be obtained if shimming of the static field is performed prior 
to each imaging session. Second, imaging at higher fields (say 4 T rather 
than 1.5 T) may not increase the acquired .sup.23 Na and .sup.39 K signals 
as much as might be expected, because in general T.sub.2 * tends to 
shorten at higher fields. The shorter T.sub.2 * may cancel the signal 
gained by going to the higher field. 
Interestingly, it was found experimentally that the .sup.23 Na T.sub.2 * in 
humans at 1.5 T was only about 10 msec, whereas a value of 25 msec was 
expected (Table 2). Unfortunately, attempts to improve T.sub.2 * by 
shimming the .sup.23 Na signal were not made because manual shimming was 
not available on our system. Recalculation of the results of FIG. 3 with 
T.sub.2 *=10 msec, however, showed similar results and the optimal 
SNR/time was still limited by the blurring constraint, Bandwidth&gt;N.sub.ro 
/T.sub.2 *. 
Gradient Performance 
The simulation results showed that although demands on gradient hardware 
for .sup.23 Na and .sup.39 K imaging can be considerable, the SNR penalty 
for poor gradient hardware is modest. This is largely because SNR/time is 
maximized for low bandwidths (FIG. 3) resulting in a low amplitude of the 
readout gradient. In addition, low bandwidth increases TR and yields a 
lower duty cycle. Nevertheless, good gradient performance is useful in 
reducing phase and slice encoding times resulting in shorter TEs, which 
can result in measurable increases in SNR due to the short T.sub.2 and 
T.sub.2 * of .sup.23 Na and .sup.39 K. 
Gradient performance may be more important for .sup.39 K than for .sup.23 
Na, largely because of the low gyromagnetic ratio of .sup.39 K (20 times 
lower than .sup.1 H, Table 2). In practice, this is not a problem because 
.sup.39 K voxels must be relatively large, which reduces gradient demands, 
due to the small MR signal. Nevertheless, it appears that even if 
additional .sup.39 K signal were available, small voxels for .sup.39 K 
imaging would be difficult to achieve due to current limitations of the 
gradient hardware. 
RF Considerations 
The studies herein used a surface RF coil for transmission and detection. 
In general, the use of surface coils results in image SNR being a function 
of distance from the coil, suggesting in the case of cardiac imaging that 
the posterior wall of the heart will be darker than the anterior. 
Interestingly, however, at least in theory this appears to be much less of 
a problem for .sup.23 Na and .sup.39 K imaging than for .sup.1 H. Because 
of the short T.sub.1 of .sup.23 Na and .sup.39 K, the MR signal for these 
nuclei is relatively independent of flip angle for high speed GRE imaging. 
For example, for FLASH imaging of .sup.23 Na (FIG. 1b), the signal 
increases from 35 to 45% of M.sub.O as flip angle decreases from 90 to 
40.degree.. Consequently, if RF power is adjusted such that a 90.degree. 
flip angle is achieved in the anterior myocardium, the signal from the 
posterior myocardium will be 28% larger (assuming a constant spin density) 
even if a flip angle of only 40.degree. is realized at the posterior wall 
(FIG. 1), potentially offsetting the signal loss due to the distance away 
from the coil. 
For human imaging, it is important to ensure that the Food and Drug 
Administration (FDA) limit for Specific Absorption Rate (SAR) of RF energy 
is not exceeded (8 Watts/kg for a localized coil). SAR is proportional to 
the square of RF power (B.sub.1) and the square of frequency (.omega.), 
.omega..sup.2 B.sub.1.sup.2 Perman, W. H., Turski, P. A., Houston, L. W., 
Glover, G. H. and Hayes, C. E. Methodology of in vivo humans sodium 
imaging at 1.5 T. Radiology 160:811-820, (1986), Vlaardingerbroek, M. T. 
and den Boer, J. A. Magnetic Resonance Imaging, Berlin:Springer-Verlag, 
1996). For a constant flip angle and square pulse duration, the required 
B.sub.1 varies inversely with .gamma., i.e. a .sup.23 Na pulse requires 
3.8 times more RF power than a similar .sup.1 H pulse. The increase in RF 
power, therefore, exactly cancels the decrease in SAR due to the lower 
.omega. of .sup.23 Na as pointed by Perman et al. Methodology of in vivo 
humans sodium imaging at 1.5 T. Radiology 160:811-820, (1986). 
Consequently, at a given field SAR is independent of nucleus type. 
Therefore, to examine the SAR for human .sup.23 Na images, the identical 
imaging pulse sequence was run at the .sup.1 H frequency. The Siemens 
software reported a localized SAR of 0.75 W/kg for the quadrature head 
coil. Doubling this value to account for the fact that the coil was not 
quadrature suggests that the localized SAR for our .sup.23 Na images was 
about 1.5 W/kg, i.e. well below the FDA limit. 
The RF coil used in the present study FIG. 6 is probably far from optimal. 
Direct application of surface coil technologies such as the use of 
distributed capacitance and a quadrature design would likely yield a 
significant increase in .sup.23 Na and .sup.39 K signal. In addition, due 
to the low resonant frequencies of these nuclei the use of superconducting 
RF coils (Black, R. D., Early, T. A., Roemer, P. B., Mueller, O. M., 
Mogro-Campero, A., Turner, L. G. and Johnson, G. A. A high-temperature 
superconducting receiver for nuclear magnetic resonance microscopy. 
Science 259:793-795, (1993)) for .sup.23 Na and .sup.39 K imaging may 
yield a considerable reduction in the noise. 
In summary, optimization of imaging parameters dramatically improves 
.sup.23 Na and .sup.39 K imaging efficiencies and may make human .sup.23 
Na imaging of the heart practical in a clinical setting. Because Na.sup.+ 
is directly involved with cellular metabolism, the potential utility of 
.sup.23 Na imaging for the non-invasive examination of pathophysiologic 
changes in humans warrants further study. 
In addition, the present invention provides a process of identifying 
regional areas of myocardial damage in vivo. The process includes the step 
of imaging the heart in vivo using .sup.23 Na or .sup.39 K magnetic 
resonance imaging wherein regions of relative high image .sup.23 Na 
intensity and regions of relative low .sup.39 K intensity indicate the 
damages regions. The damaged regions of the myocardium are non-viable 
tissue resulting from any pathological condition such as ischemia. .sup.23 
Na and .sup.39 K images are obtains using fast gradient echo imaging 
techniques optimized to maximize signal acquisition by exploiting the 
short T.sub.1 of .sup.23 Na and .sup.39 K thereby reducing imaging times 
to a few minutes. 
Where the heart is imaged using .sup.23 Na imaging, that imaging is 
accomplished using a) an imaging time of from about 3 minutes to about 20 
minutes; b) about 16 phase encodes per cardiac cycle; c) an echo time of 
from about 2.5 milliseconds to about 6.5 milliseconds; d) a repetition 
time of from about 8 milliseconds to about 20 milliseconds; e) a signal 
averaging of from about 200 to about 300; and f) a voxel size of about 
1.0-1.5.times.2.0-3.0.times.4.0-8.0 millimeters. 
Preferably, the imaging time is from about 6 to about 15 minutes, more 
preferably from about 8 to about 12 minutes and, most preferably about 10 
or 11 minutes. The echo time is preferably from about 3.5 milliseconds to 
about 5.5 milliseconds and, more preferably about 4.5 milliseconds. The 
repetition time is preferably from about 10 milliseconds to about 15 
milliseconds and, more preferably about 13 milliseconds. The signal 
averaging is preferably from about 225 to about 275 and, more preferably 
about 256. A preferred voxel size is about 1.25.times.2.5.times.6 
millimeters. 
Using a .sup.23 Na imaging process of the present invention (See the 
Examples to follow for precise details), eighteen rabbits underwent in 
situ coronary artery occlusion and reperfusion. To examine .sup.23 Na 
image intensity in viable and non-viable regions, in situ myocardial 
infarction and reperfusion was followed by .sup.23 Na imaging of either 
isolated rabbit hearts (n=6) or in vivo rabbit hearts (n=6). To examine 
regional sodium content, .sup.23 Na MR spectroscopy was performed on 
tissue samples from all 6 in vivo rabbit hearts and an additional 6 
isolated rabbit hearts subjected to the same infarction/reperfusion 
protocol. In 4 of the 6 isolated rabbit hearts, .sup.23 Na relaxation 
times (T.sub.1 and T.sub.2) were also measured in the tissue samples. To 
evaluate image quality in both small and large animals, in vivo .sup.23 Na 
images of normal rabbits (n=6) and dogs (n=4) were acquired. All in vivo 
.sup.23 Na MRI was performed in closed-chest, cardiac-gated animals. 
Image intensity of non-viable myocardium was 42.+-.5% higher than that of 
viable myocardium in isolated hearts and 95.+-.6% higher in vivo. 
Spectroscopy results showed that non-viable tissue had on average a 
63.+-.8% increase in Na.sup.+ ! compared to viable tissue in isolated 
hearts and 142.+-.7% increase in vivo, showing that the differences in 
image intensity were due to differences in myocardial Na.sup.+ !. 
Similarly, previous studies have also shown increased tissue Na.sup.+ ! 
in infarcted, reperfused myocardium by in vitro techniques such as flame 
emission photometry. The finding that image intensity differences were 
smaller than postmortem sodium concentrations could be explained by 
differences in tissue .sup.23 Na relaxation characteristics: both T.sub.1 
and T.sub.2 in non-viable myocardium would have opposite effects on image 
intensity, with shorter T.sub.1 increasing image intensity and shorter 
T.sub.2 (along with T.sub.2 *) decreasing image intensity. Nevertheless, 
despite the fact that generated images have some T.sub.1, T.sup.2, and 
T.sub.2, weighting, the data suggest that regional differences in tissue 
sodium concentration are so large that tissue Na*! dominates image 
intensity. Partial volume effects, in which relatively large imaging 
voxels contain both non-viable and viable myocardium, is an additional 
factor which could lead to smaller differences in image intensity compared 
to differences in image intensity compared to differences in tissue Na*!. 
A voxel of normal myocardium would likely have a sodium concentration of 37 
mM, assuming 77% of the tissue is water, 75% of the water space is 
intracellular, Na*!.sub.1 =15 mM and Na*!.sub.0 145 mM. The spectroscopy 
results for the in vivo experiments reported herein showed virtually the 
same value (38.+-.1 mM, FIG. 9). In isolated hearts, however, the value 
was higher (61.+-.2 mM). The elevation in Na*! in isolated hearts was 
likely due to edema formation, as suggested by our measurement of tissue 
water content in isolated hearts of 85.+-.0.3% compared to ca. 77% in 
vivo. Elevated intracellular Na*! in the isolated hearts may also have 
contributed. In non-viable myocardium, Na*! was 99.+-.4 mM in vivo. These 
values are close to the value one would estimate assuming all myocytes in 
the non-viable region failed to maintain a sodium concentration gradient, 
namely 112 mM (0.77*145=112, assumes 77% of tissue is water, plasma 
Na*!=145 mM). 
Increases in tissue Na* in non-viable regions, however, require sodium 
delivery via microvascular perfusion. Jennings et al. have clearly shown 
that infarcted tissue without reperfusion may take several hours for total 
tissue sodium to rise since electrolyte delivery would depend on slow ion 
diffusion. FIG. 8 shows that .sup.23 Na image intensity decreases 
(22.+-.4%, p&lt;0.05) during complete ischemia, perhaps secondary to 
decreases in vascular and/or interstitial volumes (which contain high 
Na*!) caused by reduced perfusion. Although "no-reflow" zones in the core 
of the infarct could also limit electrolyte delivery to infarcted 
myocardium, recent studies suggest that regional no reflow due to 
microvascular damage or stasis from intravascular neutrophil accumulation 
is a progressive phenomenon that develops during the perfusion period in 
areas that initially received adequate reperfusion. 
In the present study, the composite sodium signal was obtained without 
differentiation between intracellular and extracellular signals. In 
general, relating .sup.23 Na image intensity to myocardial viability is 
considerably complicated by the contribution of extracellular Na* to image 
intensity. For example, because extracellular Na*! is normally much 
greater than intracellular Na*!, even a small increase in extracellular 
volume due to edema may significantly elevate .sup.23 Na image intensity 
while intracellular Na*! remains near normal levels. However, in vivo 
elevations in myocardial image intensity of nearly 100% and postmortem 
Na*! of nearly 150%, as found in the present study, would be difficult to 
explain without postulating substantial intracellular accumulation of Na*. 
The present studies are the first to employ recently-developed fast imaging 
techniques, originally developed for proton imaging, to the sodium 
nucleus. Using this approach, .sup.23 Na imaging times were reduces to a 
few minutes with sufficient SNR to examine regional differences in 
mycardial sodium content. The main features which allowed a reduction in 
imaging time are: 1) gradient echoes; 2) fractional echoes; 3) extremely 
short TR; and 4) imaging at the Ernst angle. Although it has been 
recognized that the T.sub.1 of sodium can allow a short TR for signal 
averaging, previous studies have not attempted partial flip angle, 
gradient-echo imaging with extremely short TR. We hypothesized that signal 
could be gained by fast gradient-echo imaging since the short T.sub.1 of 
sodium would allow large tip angle excitations even for fast pulse 
repetition times. For example, consider a spoiled-GRASS (SP-GR) sequence 
whose theoretical signal intensity is given by 
##EQU11## 
where a is the flip angle. 
FIG. 10 (solid lines) shows the solution to this equation for a TR of 13 ms 
and T.sub.1 values similar to those of protons and sodium at 4.7 T (1400 
and 30 ms, respectively). Note that at the peak of the sodium curve the 
gradient-echo sequence collects 48% of the total theoretical .sup.23 Na 
signal every 13 ms, compared to only 8% for .sup.1 H, a 6-fold increase. 
Computer stimulation of the phenomenological Bloch equation for 
steady-state coherent sequences such as GRASS showed a similar increase in 
signal (FIG. 10, dashed lines). Furthermore, compared to classic spin-echo 
.sup.23 Na imaging with one phase encode per cardiac cycle, gradient-echo 
techniques allow acquisition of multiple phase encodes per cardiac cycle 
(16 for in vivo rabbits) and therefore significantly improve the time 
efficiency of date collection. 
To be clinically useful, it is necessary to acquire sodium images of the 
heart with voxel dimensions a few millimeters on each side and imaging 
times of a few minutes. Superficially, these requirements would appear 
very difficult to meet in light of the fact that the sodium MR signals is 
approximately 10,000 times smaller than that of protons. The result of the 
studies disclosed herein show that the combination of approaches disclosed 
herein results in an increase in signal sufficient to achieve the 
requirements for clinical sodium imaging. 
The image signal was increased by working at higher field strength (4.7 T) 
than conventional scanners (1.5 T). If it is assumed that noise is 
dominated by losses in the RF receiver coil, then SNR increases with 
frequency to the 7/4 power. If noise is dominated by sample losses then 
SNR increases only linearly. Assuming an intermediate frequency dependence 
of 3/2 power, SNR is 5-fold higher at 4.7 T than at 1.5 T. Second, voxel 
volume was at least 15-fold higher than is routinely used with proton 
imaging, corresponding to a 15-fold increase in signal. Third, the signal 
averaged 256 echoes. Since SNR varies with the square root of the number 
of averages, this resulted in an additional 16-fold increase in signal. 
Finally, the addition of a 6-fold increase in signal due to the use of 
fast imaging techniques applied to the sodium nucleus (FIG. 10), resulted 
in an improved SNR nearly 4 orders of magnitude 
(5.times.15.times.16.times.6=7200). 
As described in detail hereinafter in the Examples, excellent results were 
obtained in in vivo dog .sup.23 Na imaging experiments using a surface 
coil which was too large for the dog but reasonable for humans (15 cm), 
voxel sizes similar to those which might be useful clinically 
(3.times.6.times.25 mm), and acquired .sup.23 Na images in 4 minutes. 
Image SNR was similar to routine clinical proton images (20.+-.3 in 
anterior myocardium), strongly suggesting that existing high field 
(.gtoreq.4 T), whole-body magnets could be used to produce .sup.23 Na MR 
images of the human heart with modest trade-offs in imaging time (minutes) 
and spatial resolution (voxel dimensions .sup.3 O15=2.5 times larger than 
protons). 
The Examples to follow illustrate preferred embodiments of the present 
invention and are not limiting of the specification or claims in any way. 
EXAMPLE 1 
Human Imaging 
To allow transmission and reception of RF energy at the .sup.23 Na 
frequency (16.8 Mhz), a second RF transmit/receive subsystem was 
constructed and interfaced to a 1.5 T Siemens Vision. The second RF 
channel consisted of a frequency synthesizer, broadband RF 
transmitter/receiver, transmit/receive switch, and broadband pre-amp 
(TecMag, Houston, Tex.), as well as home-built electronics used to 
interface the Siemens timing signals and RF phase modulation to the 
transmitter/receiver, and a 1.2 kW broadband RF amplifier (Henry Radio, 
Palo Alto, Calif.). 
A home-built 16 cm diameter .sup.23 Na surface coil was used for imaging. 
The circuit diagram for the .sup.23 Na coil is shown in FIG. x. The 
.sup.23 Na coil electronics were contained within a specially-designed 
Plexiglas housing to protect the volunteer. To allow examination of the 
location of the .sup.23 Na RF coil with respect to the heart in .sup.1 H 
scout images, a small chamber was machined into the center of the 
Plexiglas housing and filled with a solution of saline and 10 mM Gd-DTPA. 
The Plexiglas housing also contained three 28" long plastic rods to allow 
adjustment of the tune, match, and balance capacitors of the coil in the 
magnet without moving the volunteer. Tuning and matching of the coil was 
verified prior to .sup.23 Na data acquisition using a portable RF sweeper 
(Model 405NV, Morris Instruments, Gloucester, Ontario, Canada). 
Four volunteers were studied. The .sup.23 Na RF coil was first centered on 
the scan table. Each volunteer was then placed prone over the coil and the 
volunteer was moved into the magnet. To locate the heart exactly over the 
center of the .sup.23 Na coil, .sup.1 H scout images were acquired using 
the body RF coil. The location of the Plexiglas chamber containing the 
saline/Gd-DTPA solution was examined, and the volunteer was asked to move. 
The double-oblique plane corresponding to the short-axis of the heart was 
then established, and a final series of cardiac gated, high-resolution 
short-axis .sup.1 H images were acquired at six base-apex slice locations 
encompassing the heart for later comparison to the .sup.23 Na images. 
The Siemens .sup.1 H RF channel was then disabled and the second channel RF 
electronics were used to acquire .sup.23 Na images. An FID was acquired to 
verify that the system was on resonance and used to determine the global 
T.sub.2 * of the sodium signal for use in selecting the receiver 
bandwidth. Next, the RF power level was examined by acquiring a projection 
image of sodium with the readout direction perpendicular to the plane of 
the coil. The pulse width was varied until the signal from the heart was 
maximized. Then, a 3D GRE pulse sequence, identical to that used in the 
numerical simulations, was used to acquire the .sup.23 Na images with the 
RF power set to the optimal value determined from the projection images. 
The .sup.23 Na image planes were prescribed such that the slices of the 3D 
acquisition corresponded to the short-axis locations of the .sup.1 H scout 
images. By encoding the 3D .sup.23 Na images in the short-axis 
orientation, in-plane resolution could be improved at the expense of slice 
thickness. Imaging parameters used for human .sup.23 Na imaging at 1.5 T 
were: imaging time 15 minutes; no gating was used; TR 20 ms; TE 3.6 ms; 
typical RF pulse width 1500 .mu.s; receiver bandwidth 6,250 (.+-.3,125) 
Hz; N.sub.AVG =50; matrix size=64.times.64.times.14; voxel 
size=6.times.6.times.12 mm. 
FIG. 5 shows three short axis .sup.23 Na images of a human heart acquired 
at 1.5 T. The 3D dataset from which these images were extracted was 
acquired in 15 min. The .sup.1 H images at the same location are shown for 
comparison. The .sup.23 Na image intensity is elevated in the cavities 
(blood pool) and the myocardium is clearly visible. The .sup.23 Na SNR's 
for the images shown in FIG. 5 were 11 in the left ventricular cavity and 
7 in the myocardium (septum). Similar results were obtained in each of the 
three other volunteers. 
EXAMPLE 2 
General Methods for Viability Studies 
MR Imaging and Experimental Protocol 
All images were acquired on a GE/Bruker 4.7 Tesla Omega system using a 
gradient-echo pulse sequence utilizing basic features of gradient-recalled 
acquisition in the steady state (GRASS). For isolated hearts, the sequence 
was then continuously. For in vivo imaging, cardiac-gated, segmented 
k-space data acquisition was used. To decrease .sup.23 Na imaging times, 
half period sinusoid gradients were used for many of the gradient wave 
forms including the slice select gradient. In addition, the slice refocus, 
phase encode, and readout prephaser gradient lobes were chosen to overlap 
completely, and had a minimum duration determined by the maximum gradient 
strength and the lobe that required the greatest area. Partial-echo 
acquisition was employed to further reduce TR and TE. Different gradient 
sets were use for rabbit imaging. For rabbits, the maximum gradient slew 
rates and amplitudes were 19.5 Gauss/cm/msec and 3.9 Gauss/cm, 
respectively. For dogs, the corresponding values were Gauss/cm/msec and 
1.2 Gauss/cm. 
Image Analysis 
Isolated Hearts 
Since the epicardial marker had guided selection of both the TTC and MR 
slice, spatial correlation of the tow images was undertaken. For each 
heart, the left ventricle (LV) on the digitized TTC stained image was 
traced by two independent observers using the software package NIH Image 
on a Macintosh Quadra. The non-viable TTC negative region was also traced. 
These outlines were superimposed over the MR image which was scaled and 
rotated appropriately to match the LV borders. The TTC negative outline 
was the used to draw a comparable re-lon on the MR image. In all cases the 
reason of altered signal intensity on the MR image was similar in size and 
location to the region of abnormal TTC staining (see FIG. 7). However, 
since the LV borders on the MR image were not identical to the TTC image 
and did not perfectly overlay, observers were instructed to include 
myocardial regions with obviously altered signal intensities. 
Regions-of-interest (ROIs) were also selected from remote, viable regions 
of myocardium. Signal intensities were normalized to the saline standard 
and averaged for the two observers. 
In-Vivo Hearts 
For in vivo hearts subject to infraction, ROIs were placed over the 
infarcted territory (identified by the external marker and postmortem TTC 
staining) and an adjacent, viable region. In normal animals, ROIs were 
placed in the anterior myocardium, left ventricular cavity, and posterior 
myocardium to calculate signal-to-noise ratios (SNR) at these locations. 
All Hearts 
SNRs were determined using Henkelman's method for magnitude images. 
MR Spectroscopy 
Tissue samples (350 to 750 mg) were taken from non-viable and viable 
regions. The non-viable region, distal to the coronary occlusion site, was 
easily identified by discoloration and the presence of intramyocardial 
hemorrhage. The tissue samples were blotted dry to remove surface 
contamination. The circumferential margins of the samples were trimmed 2 
mm in case capillary action may have removed tissue water. Samples were 
weighed and placed in sealed glass tubes. Na.sup.+ concentrations of the 
tissue samples were determined spectroscopically by comparison to the 
Na.sup.+ signal of a reference standard. The standard consisted of a 
sealed glass tube filled with 1 nil of a solution containing 109 mM 
Na.sup.+ and 43.5 mM dysprosium triethylenetetraminehexaacetic acid 14 
(Dy-TTHA). The Dy-TTHA was used to shift the .sup.23 Na peak of the 
standard such that two .sup.23 Na peaks would appear in the spectrum: one 
peak from the glass tube containing the tissue and one pak from the Class 
tube containing the Dy-standard. Care was taken to place the tissue sample 
and the adjacent Dy-standard entirely within the RF .sup.23 Na spectra 
were acquired using a 90.sup.+ pulse (45 .mu.Ls), a pre-acquisition delay 
of 58 .mu.s, a data six of 1 K, an aquision of time of 100 ms, 512 
averages, and a repetition rate of 250 ms to allow complete relaxation 
between pulses. Tissue Na.sup.+ ! was calculated as: (area under tissue 
peak/area under standard peak)*109 mm*(1 g/tissue sample weight). The 
results were express as mM Na.sup.+ and assume 100% visibility for all 
sodium signals. The spectroscopic method above .sup.23 Na was validated by 
measuring sodium concentration in five test tubes containing known 
concentrations of sodium. The r value relating known to measured Na.sup.+ 
! was 0.99. 
T1 and T2 Determination 
Relaxation times were measured for total (intracellular plus extracellular) 
myocardial sodium. T.sub.1 values were obtained using standard inversion 
recover, with inversion time ranging from 1 to 250 ms (9 data points). A 
Hahn spin-echo pulse sequence with echo times ranging from 0.5 to 40 ms 
used to measure T.sub.2 (11 data points). For both experiments there were 
64 averages with a pre-delay of 250 ms. All signals were analyzed as the 
area under the peak in the frequency domain. The T.sub.1 data were fit to 
a single exponential. The T.sub.2 data were fit to a double exponential 
using the equation. 
EQU M=M.sub.fast exp (.sup.-TE/T2fast)+M.sub.slow exp (.sup.-TE/T2slow) 
where M.sub.fast (M.sub.slow) represents the magnitude of the signal with 
time constant T21.sub.fast (T.sub.slow) and where TE is the echo time. The 
sum of the coefficients M.sub.fast and where TE is the echo time. The sum 
of the coefficients M.sub.fast and M.sub.slow was constrained to equal 
1.0. 
Tissue Water Content 
In the isolated hears, we/dry weights of viable and non-viable myocardium 
were measured by desiccation in a heating oven at 50.degree. C. for at 
least 36 hours to determine if differences in tissue sodium content could 
be explained by edemana 
Statistical Analysis 
All results were expressed as mean.+-.SEM. Differences between viable and 
non-viable myocardium in image intensity, sodium content, tissue water 
content, and relaxation characteristics were assessed using t test. The 
hypothesis that image intensities in the same region of the same heart 
varied before, during, and after coronary artery occlusion was assessed 
using repeated-measures ANOVA. Comparison between MRI and MRS results was 
accomplished using an unpaired t test for isolated hears and a paired t 
test for in vivo hearts. Values of p&lt;0.05 were considered significant. 
EXAMPLE 3 
In situ and in vivo .sup.23 Na Imaging 
In situ procedures--New Zealand White rabbits (3.5-4.0 kg) were 
anesthetized with intravenous sodium pentobarbital (ca. 27 mg/kg), 
intubated, and mechanically ventilated. A median sternotomy was performed, 
and a reversible snare ligature was placed around an anterior branch of he 
left coronary artery. After 40 minutes of in situ occlusion followed by 60 
minutes of reperfusion, the hearts were rapidly excised and retrogradely 
perfused with cardioplegic solution at room temperature. An epicardial 
marker (2-mm. Diameter polyethylene tube filled with saline) was attached 
to the right ventricle (RV) at the same base-to-apex level as the infarct 
territory. Pressure was adjusted at the beginning of the experiment of 
obtain a flow of 10 ml/min (1.0-1.5 ml/min/g, measured with an in-line 
electromagnetic flowmeter, model 1401, Skalar Medical) and then held 
constant. Typical perfusion pressures were 35-45 mmHg. The perfusate was 
not recirculated. Perfusate composition was (in mM): Na.sup.+ 120, K.sup.+ 
16, Mg.sup.2+ 16, Cl 160, HCO.sub.3. The perfusate was equilibrated with 
95% O.sub.2, 5% CO.sub.2, to maintain pH 7.4-7.53. Previous studies show 
that hearts isolated in this manner remain viable. The hearts were hung 
vertically in a 30-mm diameter radio frequency (RF) volume coil and placed 
in the magnet. 
A test tube filled with normal saline (Na.sup.+ !=154 mM) was placed 
adjacent to the heart for signal calibration. Left ventricular (LV) short 
axis .sup.23 Na images were acquired using the epicardial marker to locate 
the appropriate slice. Imaging parameters were: imaging time 7.1 minutes; 
TE 4.6 ms; TR 13 ms; N.sub.AVG =256; matrix size=256.times.128; voxel 
size=0.6.times.1.2.times.4.5 nun. Imaging was performed at the Ernst angle 
which was determined empirically. 
Following MR imaging, the short axis slice identified by the epicardial 
marker (ca. 4 mm thick) was incubated in a 1% TTC (triphenyltetrazolium 
chloride) solution at 37.degree. C.-40.degree. C. for 15 minutes. Since 
TTC forms a red precipitate in the presence of intact dehydrogenase enzyme 
systems and reducing coenzymes, viable myocardium stains brick red whereas 
necrotic areas fail to stain. The TTC stained myocardial slice was 
photographed, and the resultant 35 mm slides were digitally scanned for 
subsequent analysis. 
In- Vivo Procedures--New Zeland White Rabbit were anesthetized with 
intramuscular ketamine/xylazine (50 mg/kg and 2.5 mg/kg respectively), 
intubated, and mechanically ventilated. A catheter was placed in the 
femoral artery to monitor systemic pressure. A left thoracotomy was 
performed at the fifth intercostal space. A deflated 2-mm angioplasty 
balloon catheter was loosely asutured around an anterior branch of the 
left coronary artery. An epicardial marker filed with saline was placed 
over the territory perfused by the artery, and a catheter was placed in 
the left atrium for injection of fluorescent microspheres (15 .mu.m, 
Molecular Probes). The chest was then closed in two layers, the rabbits 
were placed prone on a 5-cm diameter gouble-resonant .sup.23 Na-.sup.1 H 
surface RF coil, and placed in the magnet. Using this approach, coronary 
artery occlusion and reperfusion could be performed closed-chest in the 
magnet by inflation ad deflation of the ballon. 
Femoral artery pressure was used for cardiac gating. Double-oblique, short 
or long-axis .sup.1 H images were first acquired using the epicardial 
marker to identify the to-be-infracted territory. The RF coil was then 
tuned to the .sup.23 Na frequency and a control .sup.23 Na frequency and a 
control .sup.23 Na image was acquired at the same location. .sup.23 Na 
imaging parameters were: imaging time 11 minutes; 16 phase encodes/cardiac 
cycle (gated to end-diastole); TE 4.6 ms; TR 13 ms; NAIG+256: matrix 
size+256.times.128 voxel size 1.25.times.2.5.times.6 mm.-Heart rate in 
these anesthetized rabbits was approximately 180 BPM. A control set of 
microspheres was injected into the left atrium. The balloon catheter was 
then inflated to produce coronary artery occlusion for 40 minutes, a 
second set of microspheres injected, and another .sup.23 Na image was 
acquired. The balloon was then deflated to allow reperfusion, a third set 
of microspheres injected, and another .sup.23 Na image was acquired. After 
approximately 60 minutes of reperfusion, a final set of microspheres was 
injected and a final .sup.23 Na image acquired. The hears were then 
removed and sectioned at the level of the epicardial marker. One side of 
the heart was stained with TTC to verify the location and extent of 
infarction. The other side of the heart was used to obtain tissue samples 
from infarcted and normal regions for spectroscopic analysis of sodium 
content (see MR Spectroscopy) and for Microsphere flow determination. 
Normal Animals--To explore the clinical potential of sodium imaging, we 
acquired in .sup.23 Na images in normal rabbits (n=6) and dogs (n=4). 
Normal rabbits were imaged using the same methodology used for animals 
subject to infaction. Normal mongrel dogs (20-25 kg) were anesthetized 
with sodium pentobarbital and intubated. After a femoral catheter was 
inserted for cardiac gating, the animals were placed in the left 
anticubital position on a 15-cm diameter double resonant .sup.23 Na-.sup.+ 
H surface coil and placed in the magnet. The same pulse sequence was used 
for dogs. 
Double-oblique short axis .sup.1 H images were first acquired and then 
followed by .sup.23 Na images at the same location. Dog imaging parameters 
were: imaging time 4 minutes; 32 phase encodes/cardiac cycle (gated to 
end-diastole); TE 3.9 ms; TR 8.1 ms; N.sub.AVG =128; matrix=256.times.128; 
voxel size=3.times.6.times.25 mm. Heart rate in these anesthetized dogs 
was approximately 120 BPM. 
The results of these studies are summarized in FIGS. 7-10. FIG. 7 shows 
microsphere flow for the in vivo animals, demonstrating successful 
coronary artery occlusion and reperfusion closed chest in the magnet. In 
all cases, the .sup.23 Na signal was sufficient to generate clear images 
of both left and right ventricular walls. As expected, the structures with 
the highest sodium concentration (i.e. the ventricular cavity filled with 
saline perfusate or blood) had the highest image intensity. Viable 
myocardium had less signal in comparison, being 48.+-.5% of saline 
(isolated hearts) and 50.+-.3% of blood (in vivo), consistent with active 
transport of Na.sup.+ out of the myocyte. In non-viable regions, 
identified by the lack of TTC staining, Na.sup.+ image intensity was 
greater than in viable regions both in isolated hearts and in vivo, 
consistent with intracellular accumulation of Na.sup.+. 
In three of the six in vivo animals, Na.sup.+ images of the 
to-be-infarcted territory were acquired before, during, and after coronary 
artery occlusion. In the remaining three animals, only post-reperfusion 
Na.sup.+ image data were acquired. Before occlusion, image intensity in 
the to-be-infarcted territory was similar to adjacent, viable regions. 
During occlusion, Na.sup.+ image intensity within the territory decreased 
by 24% in this animal. Following 1 hour of reperfusion, image intensity 
within the territory increased by 96%. 
FIG. 8 summarizes the image intensity results. In isolated hearts, image 
intensity was 42.+-.5% greater in non-viable vs. Viable myocardium 
(p&lt;0.02). For in the in vivo hearts, the elevation was 95.+-.6% (p&lt;0.001). 
For the three hearts in which Na.sup.+ images were acquired before, 
during, and after coronary artery occlusion, Na.sup.+ image intensity 
fell by 22.+-.4% (p&lt;0.05) during occlusion and rose by 104.+-.8% (p&lt;0.001) 
following one hour of reperfusion. 
FIG. 9 shows the composite sodium concentration (intracellular plus 
extracellular) of non-viable myocardium compared with viable myocardium 
for the isolated and in vivo hearts. Sodium content was significant higher 
in non-viable tissue (isolated: 99.+-.4 mM, in vivo: 91.+-.2 mM) compared 
to viable tissue (isolated; 61.+-.2 mM, in vivo: 38.+-.1 mM, p&lt;0.001 for 
both). The elevation in sodium concentration between non-viable and viable 
myocardium measured by spectroscopy (isolated: 63.+-.8%, in vivo: 
142.+-.7%) was larger than the elevation in image intensity measured by 
MRI (isolated: 42.+-.5%, in vivo 95.+-.6%, p&lt;0.05 for both). 
All T.sup.1 relaxation data were well characterized by a single exceptional 
decay. The measurement T.sup.1 and T.sup.2 values of the Dy-standard were 
nearly the same with each experiment demonstrating the reproducibility of 
the measurements. The mean T.sup.2 of the Dy-standard was near the means 
T, value although it was slightly decreased (28.2.+-.0.3 ms vs. 
29.2.+-.0.3 ms; p&lt;0.001). Likewise, the slow component of T.sub.2 for both 
non-viable (21.9.+-.1.2 ms) and viable (31.5.+-.0.8 ms) tissue approached 
the T.sub.1 values (26.2.+-.1.5 ms and 34.2.+-.0.9 ms, respectively) 
although they were consistently less (p&lt;0.01 for both). The T.sub.1 of 
viable tissue (34.2.+-.0.9 ms); p&lt;0.005. Similarly, both the fast and slow 
component of non-viable tissue T.sub.2 were found to be shorter than 
viable tissue T.sub.2, although only difference in T.sub.2slow reached 
statistical significance (T.sub.2fast : 2.2.+-.0.2 ms vs. 3.6.+-.0.6 ms, 
p=NS; .sub.T2slow : 21.9.+-.1.2 ms vs, 31.5.+-.0.8 ms; p&lt;0.001). The 
magnitude of the fast component as a percentage of the total signal was 
not significantly different between non-viable and viable tissue (26.+-.5% 
vs. CD 22.+-.1%). 
In the isolated hearts, there were not significant differences in tissue 
water content between non-viable and viable myocardium (84.+-.0.3 vs. 
85.+-.0.3 percent water by weight respectively, p=NS). 
In-vivo sodium images of normal rabbits were similar to those in animals 
subject to infarction. Since a surface coil was used for in vivo imaging, 
the anterior LV myocardium had higher SNR than the posterior myocardium 
(rabbits: anterior 12.+-.1, posterior 8.+-.1; dogs: anterior 20.+-.3, 
posterior 11.+-.2). Compared to rabbits, we obtained in vivo sodium images 
in dog's with a shorter imaging time (4 minutes vs 11 minutes) and almost 
twice the SNR, most likely because larger animals allow the use of larger 
voxels. The imaging parameters in dogs were chosen to allow estimations of 
voxel sizes, imaging times, and image quality in humans. 
EXAMPLE 4 
.sup.23 Na and .sup.39 K Imaging Studies 
To evaluate in vivo .sup.23 Na and .sup.39 K imaging, three surface RF 
coils were constructed: a 5 cm diameter .sup.23 Na-.sup.1 H coil for 
rabbit imaging at 4.7 T; a 4 cm diameter .sup.39 K-.sup.1 H coil imaging 
at 4.7 T, and a 15 cm diameter .sup.23 Na coil for human imaging at 1.5 T. 
To evaluate the performance of each coil, the unloaded and A loaded "Q" 
values were evaluated at .sup.23 Na or .sup.39 K frequency on a 
Hewlett-Packard Spectrum Analyzer. 
New Zealand White rabbits were anesthetized with intramuscular 
ketamin/xylazine (50 mg/kg and 2.5 mg/kg respectively), intubated, and 
mechanically ventilated. Anesthesia was maintained by isoflurane using a 
pediatric anesthesia machine. A catheter was placed in the femoral artery 
for cardiac gating and to monitor systemic pressure. All images were 
acquired on a GE/Bruker 4.7 Tesla Omega system using a 3D gradient-echo 
pulse sequence. For .sup.23 Na imaging, the rabbits were placed prone on a 
5-cm diameter dual-tuned .sup.23 Na .sup.1 H surface RF coil (one peak at 
52.9 Mhz and one at 200), and placed in the magnet. For .sup.39 K Imaging, 
a 4-cm diameter dual-tuned .sup.39 K-.sup.1 H surface RF coil was used 
(one peat at 9.34 Mhz and one at 200). Cardiac-gating was used in all 
cases, and approximately 50% of the R--R interval was used to acquire the 
MR data, gated to end-diastole. 
3D .sup.1 H scout images were first acquired to determine if the RF coil 
was exactly over the heart. If not, the rabbit was moved and .sup.1 H 
image was then acquired to record the location and geometry of the heart 
for later comparison to the .sup.23 Na and .sup.39 K imaging. .sup.1 H 
imaging parameters were: imaging time 11 min; 4 phase encodes/cardiac 
cycle (gated to end-diastole); TE 2.5 ms; TR 50 ms; N.sub.AVG =8; matrix 
size=128.times.64.times.64; voxel size=0.75.times.1.5.times.1.5 mm 
zero-filled during reconstruction to yield 0.75.times.0.75.times.0.75 mm 
resolution. 
For .sup.23 Na imaging, the RF coil was then tuned to the .sup.23 Na 
frequency without moving the rabbit, and a 3D .sup.23 Na image was 
acquired at the same location. .sup.23 Na imaging parameters were: imaging 
time 20 min; 4 phases encodes/cardiac cycle (gated to end-diastole; TE 2.5 
ms; TR 50 ms; N.sub.AVG =32; matrix size=64.times.32.times.32; voxel 
size=1.5.times.3.times.3 mm zero-filled during reconstruction to yield 
1.5.times.1.5.times.1.5 mm resolution. 
For .sup.39 K imaging, .sup.1 H scouting was performed as described above. 
The RF coil was then tuned to the .sup.39 K frequency without moving the 
rabbit, and a 3D .sup.39 K image was acquired at the same location. 
Imaging parameters for .sup.39 K imaging were: imaging time 3 hours; 4 
phase encodes/cardiac cycle (gated to end-diastole); TE 2.8 ms; TR 50 ms; 
N.sub.AVG =512; matrix size=64.times.32.times.32; voxel 
size=4.times.8.times.8 mm zero-filled during reconstruction to yield 
4.times.4.times.4 mm resolution.