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Am J Physiol Heart Circ Physiol 291: H368-H378, 2006. First published December 22, 2005; doi:10.1152/ajpheart.01048.2005
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Contactless magnetocardiographic mapping in anesthetized Wistar rats: evidence of age-related changes of cardiac electrical activity

Donatella Brisinda, Maria Emiliana Caristo, and Riccardo Fenici

Biomagnetism Center, Clinical Physiology, Catholic University of the Sacred Heart, Rome, Italy

Submitted 4 October 2005 ; accepted in final form 19 December 2005


   ABSTRACT
TOP
 ABSTRACT
METHODS
 RESULTS
 DISCUSSION
 GRANTS
 REFERENCES
 
Magnetocardiography (MCG) is the recording of the magnetic field (MF) generated by cardiac electrophysiological activity. Because it is a contactless method, MCG is ideal for noninvasive cardiac mapping of small experimental animals. The aim of this study was to assess age-related changes of cardiac intervals and ventricular repolarization (VR) maps in intact rats by means of MCG mapping. Twenty-four adult Wistar rats (12 male and 12 female) were studied, under anesthesia, with the same unshielded 36-channel MCG instrumentation used for clinical recordings. Two sets of measurements were obtained from each animal: 1) at 5 mo of age (297.5 ± 21 g body wt) and 2) at 14 mo of age (516.8 ± 180 g body wt). RR and PR intervals, QRS segment, and QTpeak, QTend, JTpeak, JTend, and Tpeak-end were measured from MCG waveforms. MCG imaging was automatically obtained as MF maps and as inverse localization of cardiac sources with equivalent current dipole and effective magnetic dipole models. After 300 s of continuous recording were averaged, the signal-to-noise ratio was adequate for study of atrial and ventricular MF maps and for three-dimensional localization of the underlying cardiac sources. Clear-cut age-related differences in VR duration were demonstrated by significantly longer QTend, JTend, and Tpeak-end in older Wistar rats. Reproducible multisite noninvasive cardiac mapping of anesthetized rats is simpler with MCG methodology than with ECG recording. In addition, MCG mapping provides new information based on quantitative analysis of MF and equivalent sources. In this study, statistically significant age-dependent variations in VR intervals were found.

magnetocardiography; cardiac mapping; ventricular repolarization; aging; sex

ACCORDING TO RECENT GUIDELINES (7, 21, 26, 27, 35, 39, 48, 60, 63, 65), several ECG indexes, such as QT duration and its dispersion, are used to identify risk of sudden death and assess potential cardiotoxicity of new drugs. The latter requires a large number of animal studies in the preclinical phase of new drug development. For noninvasive assessment of ventricular repolarization (VR) in small experimental animals, the most frequently used method is the 12-lead ECG (24). Extensive body surface potential mapping (BSPM), although more sensitive than the standard ECG for evaluation of repolarization inhomogeneity (2, 19, 38, 57, 70), is difficult in small animals, and its use is limited (8, 49).

Magnetocardiography (MCG), an easier method for simplification of noninvasive cardiac electrophysiological mapping, can be an appealing alternative to BSPM. Multichannel MCG mapping measures the magnetic fields (MF) generated by cardiac activation currents, with minimal distortion due to the shape and conductivity of the lungs and torso (20, 28). The advantages of contactless MCG mapping are as follows: 1) the ability to study conscious animals without movement artifacts (65), 2) the potential to provide diagnostic information not revealed by ECG (1, 29, 36, 37, 41, 43, 44, 51, 54, 62, 64, 69, 71, 74, 75), 3) the fixed-sensors geometry (42), which minimizes errors in localization of intracardiac sources (30) and three-dimensional electroanatomic imaging of arrhythmogenic phenomena (77), and 4) the localization of VR heterogeneities associated with areas of myocardial injury (50).

A major problem with MCG mapping of small hearts might be an unfavorable signal-to-noise ratio (S/N), because the cardiac sources are very weak. For this reason, most experimental MCG studies of small animal hearts have been performed in magnetically shielded rooms (4, 22, 23, 65, 72).

On the other hand, after recent demonstrations that MCG mapping of small intact animals is feasible with less expensive unshielded instrumentation (11, 12), more extensive use of MCG for large-scale serial investigation of animal models can be expected.

In a previous pilot study, we found that the minimum animal weight compatible with detection of atrial and ventricular MF is ~250 g (11) and that cardiac activity of young healthy rats can be magnetically studied, with resolution sufficient to detect sex-related differences in VR parameters (12, 13).

The aims of this longitudinal study, carried out at different epochs in intact Wistar rats, were 1) to assess age-related changes in cardiac intervals, 2) to quantify MCG parameters of VR, and 3) to create a normality database, which is useful for pharmacological studies and interpretation of abnormal MCG patterns in animal models of cardiomyopathy.


   METHODS
TOP
 ABSTRACT
 METHODS
RESULTS
 DISCUSSION
 GRANTS
 REFERENCES
 
Experimental Animals

Twenty-four Wistar rats (12 male and 12 female) were randomly selected from a larger population of healthy rats that were clinically monitored in the animal facility of Catholic University of Rome.

Two sets of measurements were obtained from each animal: in June 2004, at 5 mo of age [328 ± 19.5 g body wt (males) and 267 ± 22.3 g body wt (females)], and in March 2005, at 14 mo of age [680 ± 51 g body wt (males) and 346 ± 49 g body wt (females)]. The investigation was approved by the Catholic University Ethical Committee and conforms to the National Institutes of Health (NIH) Guide for the Care and Use of Laboratory Animals (NIH Publication No. 85-23, revised 1996).

MCG Mapping Systems

The 36-channel MCG mapping system (Fig. 1A; CardioMag Imaging, Schenectady, NY) features 36 DC-SQUID sensors (20) coupled to second-order axial gradiometers (19-mm pick-up coil and 55- to 70-mm baselines), with an intrinsic sensitivity of 20 fT/Formula in the frequency range of interest for clinical MCG mapping (direct current to 100 Hz) (13–15). The sensors are immersed in liquid helium inside a cylindrical cryostat that was supported by a gantry system, which allows vertical adjustment and does not interfere with the experimental protocol. With a single data acquisition, the z component (Bz) of a local MF at 36 positions in a plane (6 x 6 grid, covering a 20-cm2 area) is recorded. One reference ECG lead (lead II) was always recorded simultaneously with MCG signals and used to trigger the averaging of MCG signals. All signals were low-pass filtered with a cutoff frequency of 100 Hz and then digitally recorded at 1 kHz with a Windows NT-based acquisition system (24-bit analog-to-digital conversion with automatic electronic noise rejection). No high-pass filter was used.


Figure 1
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  		Fig. 1. Overview of 36-channel magnetocardiographic (MCG) recording system (A) and digital fluoroscopy (B), with the animal in place. C: frontal view of the rat's chest. Note definition of the cardiac silhouette. D: averaged MCG waveforms (usually recorded by 16 central sensors, left), "butterfly" superposition of MCG waveforms (middle), and typical map of the z component of the cardiac magnetic field (MF), calculated simultaneously with the peak of the QRS wave (right). Black (negative) area indicates field leaving the anterior chest wall; white (positive) area indicates field entering the body (curved dashed arrow). Step between the 2 isocontour lines is 0.1 pT.

 
The relative position of the animal with respect to the MCG sensors was defined by projection of three laser beams from fixed points of the cryostat onto the animal through a transparent Perspex grid with 2-mm lead markers corresponding to the center of each sensor.

A portable digital fluoroscopy system is available (Fig. 1B), so that, after MCG mapping, frontal and lateral fluoroscopic images can be digitally acquired and stored for multimodal integration. The projection, on the animal's fluoroscopic images, of the lead markers corresponding to the laser beams allows precise localization of the position of the heart with respect to the sensors (Fig. 1C).

Experimental Protocol

The animals were anesthetized with an intramuscular injection of ketamine (0.5 mg/kg body wt) and medetomidine (75 mg/kg body wt) to reach a level of anesthesia adequate to obtain the MCG recordings, placed supine on a small cradle, and covered with surgical clothing to avoid contact with the MCG instrumentation. The distance between the bottom of the cryostat and the anterior chest wall was constant for all animals (15 mm).

All recordings were performed in normal sinus rhythm with the animal breathing spontaneously. The typical duration of MCG mapping was 300 s. Each recording was repeated twice to test for reproducibility.

At the end of the procedure, the animals were kept under control until they regained full consciousness; then they were returned awake to the animal housing facility.

Signal Analysis

Postprocessing and MCG signal analysis was carried out with the same approach used for clinical investigation: 1) Windows NT-based CardioMag analysis (13–15) and 2) UNIX-based MCG software developed at Helsinki University of Technology (Neuromag) (52).

To eliminate the power-line noise, MCG signals and reference ECG were digitally filtered with an adaptive filter at 50-Hz frequency. No additional low-pass filter was applied.

The peaks of the R waves were automatically identified from the reference ECG and used as the time reference for signal averaging, with noisy and/or premature beats removed before averaging. A baseline was defined by automatic (or interactive) selection of two time points between the end of the T wave and the beginning of the P wave and subtracted from the averaged traces.

MF were accepted as "signals" only when their strength was at least three times that of the average noise (measured in the TP interval). MCG averaged data were analyzed as waveforms and as contour maps. An averaged MCG waveform is similar in morphology to an averaged ECG (Fig. 2).


Figure 2
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  		Fig. 2. Morphology of the MCG signal, recorded from grid position C4, is usually similar to that of ECG lead II (13). MF distribution computed simultaneously with peaks of P (Ppeak), QRS, and T (Tpeak) waves and at the end of the T wave (Tend) is typically dipolar. For explanation of polarity of MF, see Fig. 1.

 
Contour maps of MF distribution were automatically constructed by interpolation, from all averaged signals, of isofield lines at times of interest with a time resolution of 1 ms (Fig. 1D). Contour maps were also constructed after time integration of specific intervals of interest.

For measurements of the time intervals, the peak of the P wave, onset of the Q wave, J point, and peak of the T wave (Tpeak) were manually selected using a "butterfly" superposition (Fig. 1D) of all MCG signals amplified at a resolution of 2 cm/pT with a time scale of 200 mm/s.

Determination of time points from the MCG waveforms by morphological analysis of the dynamic evolution of MF maps (constructed every millisecond) helped improve the accuracy of QRS and JT interval timing, because each instant of interest was clearly identified by a change in polarity and/or morphology of the maps (Fig. 2) (13).

As in the murine ECG (24), in Wistar rats it is difficult to identify the end of ventricular depolarization and the beginning and end of VR from MCG waveforms because of the absence of a well-defined isoelectric ST segment separating the T wave from the QRS interval. Therefore, in several studies, the large and slower wave [corresponding to the b wave of Danik et al. (24)], always appreciable after the end of the QRS interval, was arbitrarily accepted as the T wave (13, 25). Consequently, the JT interval was divided into two segments: JTpeak and JTend. QTpeak, QTend, and Tpeak-end were measured accordingly (13). In measuring VR intervals from ECG lead II, the end of the T wave (Tend) was defined by the intersection of the tangent to the fastest portion of the descending limb and the baseline (10).

MCG Cardiac Intervals

RR and PR intervals, QRS segment length, and QTpeak, QTend, JTpeak, JTend, and Tpeak-end were measured. The onset and offset of each interval were visually selected from the butterfly superposition of all MCG waveforms and refined by analysis of the morphology of the MF maps. To correct for interindividual heart rate (HR) variation, VR intervals were corrected as follows: the measured values were divided by the square root of the averaged R-R interval [corrected value = measured value (ms)/Formula (s)].

Quantitative Analysis of VR Maps

The following parameters, equivalent to those typically used in clinical studies to detect VR abnormalities in patients with coronary artery disease or cardiomyopathy, were also calculated from MCG maps.

JTpeak MF gradient. JTpeak MF gradient (MFG) orientation was calculated using the surface gradient method, which is based on arrow maps, according to Hänninen et al. (37).

The location and direction of the largest spatial gradient of the signal distribution in the measurement plane were computed at two time intervals: 1) the integral of the second quarter from the J point to Tpeak, representing the ST segment, and 2) Tpeak. The MF {alpha}-angle was then calculated as the angle between the direction of the largest gradient and the animal's right-left line for both intervals: JT segment and Tpeak (Fig. 3).


Figure 3
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  		Fig. 3. A: MF distribution and MF gradient (MFG) orientation (arrow maps) automatically calculated at the peak of the T wave (Tpeak, top) and during JTpeak (bottom). B: schematic representation of {alpha}-angles of JT and T waves.

 
T-wave "extrema" MF dynamics analysis. T-wave extrema MF dynamics analysis was automatically calculated in any floating 30-ms time windows during JTpeak (starting when MF strength was equal to one-third of that at Tpeak and continuing until Tpeak) as the following three parameters (15, 31, 54): 1) change in the angle between the positive pole and the negative pole, which, in humans, is abnormal if >45°; 2) distance between the positive pole and the negative pole, which, in humans, is abnormal if >20 mm; and 3) ratio of strength of the positive pole to strength of the negative pole, which, in humans, is abnormal if >0.3.

Inverse Solution and Equivalent Source Parameters

For three-dimensional localization and imaging of cardiac sources generating the MF, the inverse problem was solved with the equivalent current dipole (ECD) in a semi-infinite space with homogeneous conductance (52), with the effective magnetic dipole (61), and with current density imaging (53).

The orientation (degrees) and strength (µA) of the equivalent current sources at the peaks of the P, QRS, and T waves were calculated simultaneously.

Statistical Analysis

Values are means ± SD. Statistical analysis was performed with the paired two-tailed Student's t-test to evaluate the significance of age- and sex-related differences. P < 0.05 was considered significant.


   RESULTS
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 ABSTRACT
 METHODS
 RESULTS
DISCUSSION
 GRANTS
 REFERENCES
 
MCG Cardiac Intervals

Average values of all cardiac intervals, measured at 5 and 14 mo of age, are summarized in Tables 1 and 2.


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  		Table 1. MCG-measured uncorrected time intervals

 

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  		Table 2. MCG-measured HR-corrected time intervals

 
After digital adaptive filtering of power-line interference (50 Hz), S/N of beat-to-beat MCG signals was >3 at the peak of QRS (QRSpeak) in all animals. After 300 s of recording were averaged, S/N improved as the square root of the number of averaged beats, thus a factor of 10–15, as a function of individual HR. S/N was 3–4 for the P wave, 10–12 for QRS, and 5–6 for JT-T, except in two 5-mo-old animals. Thus the quality of averaged MCG waveforms was comparable with that of ECG recordings (Fig. 2) and adequate for accurate measurements of all cardiac intervals in 22 of 24 young animals. In two rats (1 male and 1 female), S/N was <3 during the JT-T interval and, thus, was not acceptable for precise definition of VR intervals; therefore, these two animals were excluded from the statistical analysis of cardiac intervals. In the remaining 22 animals, average R-R, PR, and QRS intervals were significantly shorter (P < 0.01) at 5 than at 14 mo of age, without significant sex-related differences.

A statistically significant (P < 0.001) age-related prolongation of uncorrected QTpeak, QTend, JTpeak, JTend, and Tpeak-end was found when data from 5- and 14-mo-old animals were compared (Table 1). However, only QTend (P < 0.001), JTend (P < 0.01), and Tpeak-end (P < 0.001) remained significantly longer at 14 mo of age, if repolarization intervals were corrected for the HR, according to the Bazett's formula (Table 2).

When we accounted for sex-related differences, at 5 mo of age, uncorrected QTend (P < 0.02), JTend (P = 0.01), and Tpeak-end (P < 0.05) were significantly longer in female than in male rats (Table 1). However, HR-corrected values did not differ significantly at this age: HR was significantly slower in female than in male rats (P < 0.02; Table 2).

At 14 mo of age, no difference in uncorrected VR intervals was found among female and male rats. On the contrary, HR-corrected QTend, JTend, and Tpeak-end were significantly longer in male rats (P < 0.05; Table 2); HR was significantly lower in female rats at 14 mo of age as well.

Although a comparison between MCG and ECG was not an end point of this study, VR intervals were also measured from ECG lead II. In contrast to multisite MCG recordings, Tend was not easily defined from a single ECG lead at 5 mo of age, because the amplitude of the T wave was very low. The average ECG values for QTend, JTend, and Tpeak-end were significantly shorter than the corresponding MCG measurements at 5 and 14 mo of age (Tables 1 and 2).

Quantitative Analysis of VR Maps

An example of typical MF distribution and MFG orientation computed at the integral of JTpeak and at Tpeak is shown in Fig. 3.

Average values of quantitative parameters of VR calculated from MCG maps are summarized in Table 3. No significant age- or sex-related difference was found between the two sets of recordings.


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  		Table 3. Quantitative analysis of VR maps

 
Inverse Solution and Equivalent Source Parameters

Inverse localization of cardiac sources and a quantitative estimate of the ECD strength and orientation were computed simultaneously with Ppeak, QRSpeak, and Tpeak.

Although a wider variability in ECD orientation was observed for the P and T waves at 14 mo of age, no significant age-related difference in orientation of the ECD was found (Fig. 4).


Figure 4
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  		Fig. 4. Range of variability (shaded areas) of current dipole orientation measured at Ppeak, QRSpeak, and Tpeak. Data represent cumulative values of male and female rats measured at 5 and 14 mo of age.

 
The ECD moment, measured at Ppeak, QRSpeak, and Tpeak, was significantly stronger at 14 than at 5 mo of age (P < 0.05; Table 4), as expected because of the age-related increase in body weight and myocardial mass.


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  		Table 4. Equivalent current dipole strength

 
In all animals, source localization, calculated at QRSpeak, was coincident with the anatomic position of the ventricles, inferred from fluoroscopic imaging, with some differences among the different methods (Fig. 5).


Figure 5
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  		Fig. 5. Multimodal electroanatomic integration and source imaging based on MCG mapping. A: scaled MF distribution and fluoroscopy image of the animal, with lead markers (small black dots) matched with center of each MCG sensor, identified by laser beam spots. Cardiac MF spreads several centimeters from the animal's body. Inverse localization of cardiac sources, computed simultaneously with QRSpeak, properly fits into the ventricular area of the cardiac silhouette, with some differences between effective magnetic dipole (EMD, open white circle) and equivalent current dipole (ECD, white square) models. B: enlargement of A, with current density estimate at the epicardial level scaled and superimposed on the heart silhouette. Current maximum (red spot on green epicardial layer) connects EMD and ECD localizations. For explanation of polarity of MF, see Fig. 1.

 

   DISCUSSION
TOP
 ABSTRACT
 METHODS
 RESULTS
 DISCUSSION
GRANTS
 REFERENCES
 
Since the early 1970s, MCG has been proposed as the magnetic alternative to ECG and BSPM for the noninvasive study of cardiac electrophysiology (1, 20, 28, 51, 62, 69, 75). In humans, multichannel MCG systems provide simultaneous multipoint cardiac mapping, with spatial and temporal resolution adequate to reveal electrophysiological markers for risk of life-threatening arrhythmias (43, 44, 74). In experimental animal studies, only a few of the 12 ECG leads are usually recorded (9, 18, 24, 47), although high-quality multichannel BSPM in rats has been reported (3, 8, 49, 66, 67). MCG mapping has advantages over BSPM: 1) it is contactless, and 2) its sensor configuration is fixed and reproducible. Both features are obviously ideal for serial cardiac mapping of small experimental animals, in which BSPM recordings are more difficult and time consuming. Recent demonstrations that MCG mapping is feasible for conscious and unrestrained animals (65) and in unshielded laboratories (11–13) disclose a new possibility for longitudinal studies of animal models with extensive cardiac mapping, i.e., testing of new drugs or gene therapy (42).

It is well known that aging is associated with morphological and electrical remodeling of the myocardium (3, 34, 45, 58), leading to a higher incidence of life-threatening arrhythmias and sudden death (33, 55), especially in the presence of ventricular hypertrophy or other cardiomyopathy (6, 25, 26, 59, 67, 68). Therefore, the primary aim of this research was to evaluate age-related changes in cardiac intervals and VR parameters calculated from MCG maps, with the influence of sex taken into account (10, 13, 73). A secondary aim was to increase the database of MCG data of young healthy rats of both sexes to verify our previous findings (13) in a larger number of cases.

The present results confirm that the sensitivity of the unshielded MCG instrumentation was adequate to reconstruct well-defined MF distribution simultaneously with Ppeak and throughout the QT interval in the majority of the rats. In only two 5-mo-old animals was the strength of MF too weak to ensure a reliable quantitative analysis of VR parameters.

The morphological similarity between ECG lead II and MCG position C4 (Fig. 2) was confirmed in all healthy rats and is a useful landmark to ensure an appropriate positioning of the heart approximately in the center of the sensor array. This is relevant for the reproducibility of MCG measurements, especially if longitudinal studies imply serial cardiac mapping of the same animal. Because of the spine's curvature, when the animal is placed supine, if the head is turned to the right, the heart could be displaced to the right as well. Such an anomaly might introduce significant errors in a quantitative estimate of MF strength and orientation.

Absolute values of cardiac intervals were consistent with those reported in previous ECG and MCG studies (6, 13, 24, 25). Clear-cut age-related changes in the duration of PR, QRS, QTpeak, QTend, JTpeak, JTend, and Tpeak-end were found. In particular, measurements at 5 and 14 mo of age demonstrated a statistically significant prolongation of all uncorrected VR intervals. However, after correction for HR, only the age-related prolongation of QTend, JTend, and Tpeak-end was still significant (Tables 1 and 2). This could suggest that, in rats, age-related QT prolongation might be mainly due to an increase in transmural dispersion of repolarization (5). On the other hand, the loss of significance of QTpeak and JTpeak seems to be also due to the significantly lower HR in the female rats, in contrast to the well-known human findings (17).

As reported previously (13), longer VR intervals were found in female than in male rats at 5 mo of age. On the contrary, at 14 mo of age, average uncorrected values of VR intervals were similar in male and female rats, and average HR-corrected values of QTpeak, QTend, JTpeak, JTend, and Tpeak-end were even shorter in female than in male rats (Tables 1 and 2). One of the possible explanations for this finding could be a different hormonal pattern in female rats between the first and the second measurement. At 14 mo of age, corresponding to the human age of ~50–60 yr (6), a reduction of the estrogenic pattern and an increase of the androgenic component are expected (10, 73). However, we do not have objective data supporting this hypothesis, because an evaluation of the hormonal patterns had not been planned in the design of this study.

Other VR parameters estimated with quantitative analysis of the MF distribution (Table 3) did not differ significantly with age, nor did they show sex-related differences. This suggests that such parameters, which have already proven sensitive for diagnosis of myocardial ischemia and other cardiomyopathies in humans, even in the absence of ECG alteration (31, 36, 37, 54, 64, 71), could be useful to reveal VR abnormalities in animal models of cardiomyopathy, independently of sex and age.

A comparison between ECG and MCG time measurements was not an end point of this study; rather, our primary aim was to investigate the age-related variation of quantitative parameters measurable with MCG mapping. Therefore, only one ECG lead was recorded, mostly to guarantee an independent QRS trigger for the averaging of MCG signals. The finding that VR intervals measured with multisite MCG mapping were longer than those measured with a single ECG lead (Tables 1 and 2) could be due to the method used for definition of the offset of the T wave, which in some cases might disregard terminal low-amplitude components. The differences were significant only for QTend, JTend, and Tpeak-end. Alternatively, MCG mapping could have revealed a physiological VR dispersion that was undetectable with ECG, unless the recording was done with the 12 standard leads.

A typical difficulty in analyzing a "murine" ECG is defining where VR starts and ends. The results of this work confirm that the sensitivity of the unshielded MCG setup is adequate to study the part of VR corresponding to the "b wave" described by Danik et al. (24). On the contrary, no organized MF distribution was appreciable after the end of that early VR wave. To better clarify the accuracy of the MCG estimate of VR and its dispersion, a comparison with multiple simultaneous action potential recordings is needed. Optical mapping would be the ideal method to attempt the most accurate validation (7). However, to the best of our knowledge, this is impossible, because instrumentations for "in vivo" optical mapping, specifically designed to avoid magnetic artifacts during MCG recordings, do not exist. Alternatively, simultaneous recording of MCG and monophasic action potential (MAP), similar to the method used to interpret the murine ECG (24, 47), can be attempted (32, 40). In a preliminary case, an MCG estimate of QTend corresponded to ~75–80% of the duration of a single epicardial ventricular MAP recorded with a percutaneous amagnetic catheter (Fig. 6). Thus an MCG quantitative estimate of VR in rats might be sensitive to the study of transient outward current, which is the dominant current of the early repolarization phase (47), but not to the terminal part of phase 3 of repolarization. A refinement of a percutaneous method for multiple MAP recording with a single amagnetic catheter will provide a direct in vivo comparison between the dispersion of action potential duration and the MCG estimate of VR (unpublished observations).


Figure 6
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  		Fig. 6. Simultaneous MCG mapping and epicardial (Epi) ventricular monophasic action potential (VMAP) recording using the amagnetic catheter technique (32). A: epicardial VMAP. Although it is not of ideal quality (noisy and with a biphasic phase 0), epicardial VMAP has a typical "spike-and-dome" appearance (47). Note time correlation between Tpeak and the peak of the dome in MCG recording. MCG estimate of the end of the QT interval corresponds to ~80% of epicardial VMAP duration. B: frontal fluoroscopic image. Open arrow, tip of amagnetic catheter, positioned at the ventricular epicardium. Arrow maps show MF distribution and corresponding bidimensional current reconstruction, computed at Tpeak and Tend. Maximum of current is indicated by white spot on arrow maps.

 
One advantage of MCG mapping over surface potential recordings is that the cardiac MF is directly sensitive to the cellular activation currents (or "primary sources"), with minimal influence of the volume currents (or "secondary sources") flowing outside the heart (76). Thus the activation currents can be quantitatively estimated from MCG data through the solution of the inverse problem with use of different models, such as the single ECD model (30, 52), the calculation of bidimensional pseudocurrent maps (42), or the three-dimensional reconstruction of current density (50, 53).

In this study, we focused only on the quantitative estimate of the strength of the ECD moment (Table 4), although three-dimensional current density at the epicardial level was also calculated for source imaging (Fig. 5B). Atrial and ventricular dipole strengths were weaker by a factor of ~5 than those described in our previous study (13). This was due to a change in the measurement setup. In the present study, the rats were placed supine, as in the previous study, but in a new cradle, resulting in an increase in the distance between the sensors and the anterior chest wall of the animal of ~12 mm. When the absolute values of MF were corrected for the difference in distance between the sensors and the animal, the strength of the ECD moment was equivalent to that of the previous study.

A statistically significant age-related increment in strength of the ECD calculated at Ppeak, Rpeak, and Tpeak was found between measurements at 5 and 14 mo of age, consistent with the increase in body weight, which was more evident in the male rats.

A quantitative estimate of the accuracy of source localization was not a primary aim of this study. However, visual analysis confirmed that the inverse localization of ventricular sources simultaneously with QRSpeak was properly centered in the anatomic position of the ventricles inferred by biplane fluoroscopic imaging, with all models used (Fig. 5A). No attempt was made to evaluate systematically the uncertainty of localization. However, because the average ventricular dimension was on the order of ~10 mm, the uncertainty of localization was at its worst within that range. A better estimate of localization accuracy is expected with application of the amagnetic catheter method (28) also in animal studies.

Study Limitations

Experimental MCG recordings were carried out in an unshielded noisy environment and with a system designed for MCG mapping of the much larger human heart; thus the diameter of each sensor's pickup coil (19 mm) was almost as large as the diameter of the rat's heart. This could somehow reduce spatial discrimination of weak cardiac sources.

MCG mapping has been shown to be feasible in awake and unrestrained nonaggressive rodents, such as guinea pigs and rabbits (65), but not in aggressive rats. Thus, in our study, anesthesia was needed.

In terms of quality, beat-to-beat MCG signals in our study were noisier than previously reported surface and telemetric electric recordings (3, 24, 49, 56, 66, 68). Thus to obtain comparable signal resolution, it was necessary to average 300 s of continuous recording. This limits the possibility of quantitative analysis of transient events, such as unsustained arrhythmias or rate-dependent changes of VR. However, after 300 s of continuous recording were averaged, MCG resolution was sufficient to reveal reproducible age- and sex-related differences in VR intervals, with sensitivity adequate to study the major repolarization wave, usually described as the b wave (24) or J wave (47). The terminal part of VR [i.e., the "c wave" sometimes observed in the rat ECG (24)] cannot be detected with an unshielded instrument but could probably be investigated with high-resolution recordings in magnetically shielded rooms with high-end multichannel systems (42, 65). To improve the sensitivity of the unshielded MCG method, a better S/N is needed to attain a signal resolution comparable with that obtained with heavy electromagnetic shielding (65). From experience with unshielded fetal MCG (16), it is predictable that a significant improvement in signal quality can be achieved with MCG signal processing based on the independent-component analysis (46).

In conclusion, this first longitudinal study in Wistar rats confirms that multichannel MCG mapping of small experimental animals, carried out with an unshielded MCG instrument designed for clinical use, is feasible and provides reproducible quantitative estimates of VR parameters, with sensitivity sufficient to reveal age- and sex-related differences in small experimental animals.

Indeed, because it is a contactless method, MCG methodology is faster and more reproducible for multisite noninvasive surface cardiac mapping of anesthetized rats than ECG recordings. Reproducible MCG mapping can be obtained in 10 min, including postprocessing for MF reconstruction and source localization.

Although this study was carried out in anesthetized rats, MCG mapping can be repeated several times in 1 day in awake and unrestrained nonaggressive animals (65), i.e., for the study of circadian variation of the parameters of interest. This suggests that MCG methodology might have some important advantages over ECG telemetry. Recent evidence has suggested that implantation of telemetric devices might be aggravated by a high percentage of complications that significantly affect the number of implanted animals that remain available for long-term longitudinal studies (56).

As for the interpretation of murine ECG (24, 47), a better understanding of the accuracy of MCG estimates of VR inhomogeneities in rats is expected from correlative investigation carried out simultaneously with MCG mapping and multisite MAP recordings.


   GRANTS
TOP
 ABSTRACT
 METHODS
 RESULTS
 DISCUSSION
 GRANTS
REFERENCES
 
This study is based on work supported by National Science Foundation Grant 0349580. This research was partially supported by Italian Ministry of Education, University, and Research Grants 9906571299-001 and 2001064829-001 and by Sigma Tau Grant DS/2004/CR/22.


   ACKNOWLEDGMENTS
 
The authors are grateful to Drs. Carl Rosner, Alexander Bakharev, Nikolai Korsun, and Karsten Sternickel for continuous support and technical improvement of the multichannel MCG system in the framework of the ongoing scientific agreement and to Dr. Jukka Nenonen for stimulating discussion and suggestions. The authors acknowledge the outstanding and invaluable dedication of Viola Iacobini and all personnel in the animal lodging facility.

Any opinions, findings, and conclusions or recommendations expressed in this publication are those of the authors and do not necessarily reflect the views of the National Science Foundation.


   FOOTNOTES
 

Address for reprint requests and other correspondence: R. Fenici, Biomagnetism Center, Clinical Physiology, Catholic Univ. of Rome, Largo A Gemelli, 8 00168 Rome, Italy (e-mail: feniciri{at}rm.unicatt.it)

The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked "advertisement" in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.


   REFERENCES
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