RF interference suppression in a cardiac synchronization system operating in a high magnetic field NMR imaging system

Magnetic Printed

Resonance in the USA.

Imaging, Vol. 6, pp. All rights reserved.

637440,

1988 Copyright

0730-725X/88 $3.00 + .oO 0 1988 Pergamon Press plc

l Original Contribution

RF INTERFERENCE SUPPRESSION IN A CARDIAC SYNCHRONIZATION SYSTEM OPERATING IN A HIGH MAGNETIC FIELD NMR IMAGING SYSTEM

A.A. DAMJI,*

R.E. SNYDER,* D.C. ELLINGER, * F.X. WITKOWSKI~, AND P.S. ALLEN* *Department of Applied Sciences in Medicine and tDepartment of Medicine, University of Alberta, Edmonton, Alberta, Canada, T6G 2G3

An electrocardiographic (ECG) unit suitable for cardiac-synchronized nuclear magnetic resonance imaging in high magnetic fields is presented. The unit includes lossy transmission lines as ECG leads in order to supress radio frequency (RF) interference in the electrocardiogram. The unit’s immunity to RF interference is demonstrated. Keywords:

MR imaging; Cardiac synchronization;

RF interference.

By synchronizing the NMR data acquisition sequence to the cardiac cycle, one can minimize not only artifacts in the image of the heart and its surrounding vessels, but also artifacts occurring in images of other regions which arise from the motion of the heart wall or the pulsatile motion of blood.’ Lanzer et al.,* who evaluated various sources for a cardiac trigger, have shown that cardiac synchronization is optimized when the electrocardiogram (ECG) R-wave is used to generate the triggering pulses. However, several characteristics of the NMR environment degrade the ECG: the pulsed radio frequency (RF) field, the switched magnetic field gradients, ECG lead vibrations in the very large static magnetic field, and possibly the Hall effect emf due to blood flow within a static magnetic field.3 In a 40 cm bore animal imaging system, operating at a magnetic field strength of 2.35 T (equivalent to a proton resonance frequency of 100 MHz), the pulsed RF interference was found to be the overwhelming source of degradation. In comparison with clinical NMR imaging systems operating at lower field strengths, the increased RF interference observed at higher field strengths is due to the higher operating frequency as well as the higher RF pulse power available in medium bore NMR units. We briefly present

the pertinent features of a cardiac synchronization system, with particular reference to the suppression of RF interference by means of a lossy (dissipative) transmission line and to the accommodation of the triggering peculiarities of Bruker CXP spectrometers used in animal imaging systems. Initial attempts to reduce RF interference to acceptable levels by means of screened coaxial ECG leads and low-pass filters at the ECG amplifier inputs were unsuccessful. We therefore replaced these components with a non-metallic, twin-lead, lossy transmission line (Fluorosint TM 8 19 Dual Line, The Polymer Corporation, Reading, PA) and successfully removed the RF interference. To understand the reduction in interference, one might consider the two extreme cases of a discrete interference source and a uniformly distributed interference source. If the RF interference source is lumped at the input of the transmission line and reduced to a Thtvenin equivalent circuit having an open circuit voltage V, and an internal impedance ZT, with the other end of the line connected to the input impedence, Z2, of the amplifier (Fig. lA), the magnitude of the interference voltage, 1Vz2 1, at the amplifier input can be reduced to4

RECEIVED 1l/25/87; ACCEPTED 4/8/88.

versity of Alberta Hospitals Foundation. An AHFMR Medical Scientist Award (PSA) is also gratefully acknowledged. Address correspondence to A. A. Damji, Department of Applied Sciences in Medicine, University of Alberta, Edmonton, Alberta, Canada, T6G 2G3.

Acknowledgments-This work was supported by a major equipment grant from the Alberta Heritage Foundation for Medical Research (AHFMR), and operating funds from the Medical Research Council of Canada, the Alberta Cancer Board Research Initiatives Programme and the Uni631

Magnetic Resonance Imaging 0 Volume 6, Number 6, 1988

638

Table 1. Transmission line parameters for Fluorosint TM 8 19 Dual Line Parameter

zlo

Series Series Shunt Shunt

A (A) K

2

z&l

2s

(B) circuits

RF interference. (A)

modeling

Unit

1 x 106 440 x 10-9 3.6 x lo-l5 76 x 10-r’

O/m H/m S/m F/m

*A 20% uncertainty in R. was obtained from repeated measurements of this parameter. No such measurements were made for L,; Cc, or Cc.

(V/m)

ldZ2

Fig. 1. Equivalent

resistance*, R, inductance, L,, conductance, G,, capacitance, Cc

Value

capacitive susceptance dominates the shunt conductance. With this in mind, the attenuation constant, cy, and the magnitude of the propagation constant, 1y 1, may be approximated at frequency f by

Lumped input, (B) distributed input.

cr=m

(5)

and

2IzJZ2I ‘vz2’= lb’ lZoZT+ z,z, + z,2 + Z,Z,l e-as (1) I 4 I&-l e+

,

(2)

where cy is the attenuation constant, s is the transmission line length, and Z,, the characteristic impedance of the line, has a phase angle of -45” for the resonance frequencies currently encountered in NMR imaging. It should be noted that the voltage appearing at the amplifier input due to RF coupling is attenuated by a factor proportional to exp(-ols). At the other extreme, the RF coupling may be regarded as being evenly distributed along the transmission line with intensity K (V/m), as shown in Fig. 1B. With terminal impedances of Z1 and Z, attached to the line, the magnitude of the interference voltage can be written in terms of the propagation constant y as lvzzl

= IKI 5

IZZI jz*+z,/

JZ’K’ ‘r-‘1

IYI

= Jza

,

(6)

where RO is the series resistance (per conductor) and COthe shunt capacitance per unit length of transmission line (Table 1). Table 2 lists values of CYand the minimum attenuation available for a 1 m line for the cases of lumped and distributed RF coupling at various frequencies currently encountered in proton NMR imaging. It should also be pointed out that since (Yis proportional to @, RF attenuation will increase with frequency for a given length line. While the above is not intended as a rigorous description of the RF attenuation due to a lossy line under circumstances like ours, it does enable us to quantify some limits. For example, it illustrates the complete immunity that such a line would offer from an RF source coupled only to the input of the line. However, in a real experimental configuration, distributed coupling would be expected to be the important source of interference, since a significant length of the transmission line is within the RF coil of the

ly-ll

.

Table 2. Attenuation constant ((Y) and minimum attenuation for a 1 m length of the transmission line described in Table 1

(4)

The parameters of the lossy line, calculated from the manufacturer’s specifications, are presented in Table 1. For the range of proton resonance frequencies currently encountered in NMR imaging, namely 1 to 100 MHz, it should be noted that the series resistance dominates the series inductive reactance and the shunt

Frequency (MHz)

(NpTm)

Lumped (dR)

1 60 100

21.85 169.3 218.5

177.8 1459 1886

Distributed (dR) 26.8 44.6 46.8

RF Interference suppression 0 A. A. Depart ET

spectrometer and hence maximally exposed to the RF field. Nevertheless, the values presented in Table 2 for the case of distributed coupling suggest that a significant attenuation would be expected as well. Unfortunately, a rigorous quantitative measurement of the attenuation of RF interference inside the bore of a 2.35 T magnet is a major undertaking, in view of the large distance from the lossy line at which the measuring instruments would need to operate. It was not therefore performed. Estimates were made however, based upon the two lead types tested (see below, Fig. 3), and it was found that replacing the coaxial leads and low-pass filters with the twin-lead, lossy transmission line resulted in a further reduction in RF interference of > 30 dB, not inconsistent with the above analysis of distributed coupling. A synchronization/gating system was developed for a 40 cm bore animal imaging system based on a Bruker CXP spectrometer, and is illustrated in Fig. 2. Three ECG electrodes (Red DotTM Ag/AgCl with foam tape, 3M Canada Ltd., London, Ontario) were applied to the limbs of an anesthetized cat. The parallelwire transmission line, consisting of two carbon-impregnated teflon conductors enclosed in a nylon sheath, connected the active electrodes to a battery powered FET input differential amplifier placed inside the RFscreened magnet enclosure. Electrical contact to the plastic transmission line was made by wrapping a

639

AL.

short section of 30 AWG wire onto each of the exposed teflon conductors and encapsulating the wrapped segments in epoxy. An optical datalink (DMC System 5610, Dynamic Measurement Corporation, Winchester, MA) was used to transmit the amplified ECG signal to the fiber optic receiver located near the spectrometer console. A sharp 60 Hz notch filter was placed between the receiver and the trigger circuit to eliminate power-line frequency interference present in the ECG. Level triggering was used to obtain the trigger pulses, while a variable delay between the trigger and the start of the pulse sequence was incorporated into the CXP pulse program. In order to facilitate cardiac-synchronized and respiratory-gated NMR data acquisition, it was necessary to present both the cardiac trigger and the respiratory gating signals to the spectrometer. Since the CXP spectrometer is restricted to one input line, these two signals were combined by a multiplexer unit which was controlled by pulses from two software-controlled CXP spectrometer output lines. An illustration of the efficacy of the lossy transmission line is shown in Fig. 3. Fig. 3A shows the ECG recorded from a cat inside the magnet bore monitored by the circuitry shown in Fig. 2, but exposed to no RF pulses or switching gradients. The lossy transmission line was then replaced by screened leads; LC low-pass filters, which provided at least 80 dB attenuation at

ANALOG OPT I CAL DATALINK h \

/

ECG AMPLIFIER

+

TRANSHITTER

”

RECEIVER

j

60 HZ NOTCH FILTER

FROM RESP. MONITOR \1 +

TRIGGERING CIRCUITRY

MULTIPLEXER UNIT

+

TO CXP

PULSES FROM CXP Fig. 2. Schematic drawing of synchronization/gating system. Electrodes attached to the animal are connected to an ECG amplifier via the lossy transmission line described in the text. A single conductor of the same lossy line is used to connect the neutral (left leg) electrode to the common terminal of the ECG amplifier power supply. An optical datalink is used to encode and transmit the amplified ECG signal to a region removed from the magnet enclosure. The ECG amplifier and the datalink transmitter are located within the RF-screened magnet enclosure and are powered by 6 V lead/acid batteries.

640

Magnetic Resonance Imaging 0 Volume 6, Number 6, 1988 is again removed,

IS

1s

(4

,

II’

II’

Fig. 3. Electrocardiograms obtained from an anesthetized cat. (A), (C), and (D) were obtained using the lossy transmission line; (B) was obtained using screened leads and lowpass filters which provided 80 dB attenuation at 100 MHz, between the leads and the amplifier. (A) Cat in magnet bore without RF pulses or switching gradients. (B) Same as (A), but with rectangular, 125 ps, 90” RF pulses separated by 3.15 s intervals. (C) Same as (A), but with rectangular 250 ps, 180” RF pulses separated by 2.56 s intervals. (D) Multiecho imaging procedure with 2-3 mT/m, 1.5 ms rise time switching gradient pulses. Arrows in (B) and (C) indicate times of RF pulses.

100 MHz, were inserted between the leads and the amplifier. When the animal was exposed to a pulse train of rectangular, 125 ps, 90” RF pulses separated by 3.15 s intervals, the trace shown in Fig. 3B was recorded. Using screened leads and filters during multiple-echo imaging, saturation of the amplifier precluded observation of the cardiac signal for about 0.75 s and therefore prevented synchronization of the NMR sequences to successive cardiac cycles. A contrasting result was obtained when the screened leads and low-pass filters were replaced with the lossy transmission line. Fig. 3C shows the ECG recorded when the animal was exposed to a pulse train of rectangular, 250 ps, 180” RF pulses separated by 2.56 s intervals. Finally, the ECG was recorded during a multi-echo imaging procedure which incorporated both RF pulses and gradient switching (Fig. 3D). The RF interference

as in Fig. 3C, but now the interference from the gradient switching is observable. These gradients were between 2-3 mT/m in strength with a rise time of 1.5 ms. Although the gradient switching artifacts are large enough to generate false triggers, the spectrometer is insensitive to these artifacts since it polls the CXP input line only after the expiration of the relaxation period following each NMR pulse sequence. If the ECG were to be required for diagnostic purposes, then the novel idea of utilizing a time-dependent low-pass filter5 could be employed to produce ECGs of acceptable clinical quality. In addition to the gradient switching signals, the lossy transmission line also passes 60 Hz pickup. The wiring of the gradient and shim coils is the most likely means of coupling 60 Hz fields into the magnet bore. By virtue of the mismatch in the resistance of the two transmission line conductors, a 60-Hz common-mode signal at the active ECG electrodes appears as a differential signal at the amplifier inputs. It was found that a notch filter, of approximately 2 Hz bandwidth, incorporated prior to the triggering circuitry successfully eliminated 60 Hz noise. Measurements made of the linewidth of a 10 cm diameter spherical water phantom (10 Hz), with and without the presence of the equipment and materials required for ECG synchronization, showed that the ECG monitoring had no detrimental effect on the homogeneity of the magnetic field. Comparisons of images of this phantom obtained with and without the synchronization equipment present showed that the image quality had not been compromised. Finally, another positive utility of the high resistance transmission line is that patient currents are minimized and consequently patient safety is enhanced.

REFERENCES 1. Henkelman, R.M.; Bronskill, M.J. Artifacts in magnetic resonance imaging. Rev. Msg. Res. Med. 2:1-126; 1987. 2. Lanzer, P.; Botvinick, E.H.; SchiIler, N.B.; Crooks, L.E.; Arakawa, M.; Kaufman, L.; Davis, P.; Herfkens, R.; Lipton, M.J.; Higgins, B.C. Cardiac imaging using gated magnetic resonance. Radiology 150: 121-127; 1984. L.W.; Herfkens, R; Fram, 3. Dimick, R.N.; Hedlund, E.K.; Utz, J. Optimizing electrocardiograph electrode placement for cardiac-gated magnetic resonance imaging. Invest. Radiology 22: 17-22; 1987. 4. Smith Jr., A.A. Coupling of external electromagnetic fields to transmission lines. New York: John Wiley and Sons; 1977. 5. Rokey, R.; Wendt, R.E.; Johnston, D.L. Monitoring of acutely ill patients during nuclear magnetic resonance imaging: Use of a time-varying filter electrocardiographic gating device to reduce gradient artifacts. Mag. Res. Med. 6:240-245; 1988.