Impact of phase difference between cardiac systole and skeletal muscle contraction on hemodynamic response during electrically-induced muscle contractions

Journal of Electromyography and Kinesiology 20 (2010) 572–579

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Impact of phase difference between cardiac systole and skeletal muscle contraction on hemodynamic response during electrically-induced muscle contractions Tetsuya Kimura, Naoko Kameda, Toshio Moritani * Graduate School of Human and Environmental Studies, Kyoto University, Sakyo-ku, Kyoto 606-8501, Japan

a r t i c l e

i n f o

Article history: Received 16 December 2009 Received in revised form 14 March 2010 Accepted 15 March 2010

Keywords: Electrical stimulation Mechanomyogram Blood pressure Cardiac afterload

a b s t r a c t Percutaneous low-frequency electrical muscle stimulation (LF-ES) is a new alternative exercise prescription for individuals who cannot adequately perform voluntary exercise. However, substantial undesirable elevation of both systolic blood pressure (SBP) and cardiac afterload occurs during LF-ES and must be resolved. Therefore, this study examined whether or not the synchrony between cardiac systole and skeletal muscle contraction affects instantaneous blood pressure and cardiac afterload during intermittent evoked muscle contractions. In eight subjects, the quadriceps and biceps femoris muscles of each limb were simultaneously stimulated at 20 Hz with a duty cycle of 0.3 s stimulation and 0.7 s pause for 15 min. The phase difference between the ECG R-peak and the onset of muscle contraction (sc–s) was measured for all heartbeats. Then, instantaneous SBP, tension-time index (TTI), and peripheral vascular resistance (PVR) associated with each heartbeat were plotted as functions of sc–s. The results showed that SBP, TTI, and PVR were significantly lowered at positive sc–s (i.e., the moment at which a muscle contraction started during the cardiac recovery phase). These results suggest that a well-designed stimulator, one that induces muscle contractions coupled with heartbeats with appropriate phase difference, would effectively attenuate the elevation of SBP and cardiac afterload during LF-ES. Ó 2010 Elsevier Ltd. All rights reserved.

1. Introduction Percutaneous low-frequency (4–20 Hz) electrical muscle stimulation (LF-ES) has been attracting extensive attention, as it highly enhances energy expenditure (2–12 METs) without limb movement with the subject lying supine (Banerjee et al., 2005; Caulfield et al., 2004; Hamada et al., 2003, 2004). Therefore, LF-ES could become an alternative exercise prescription to prevent the development of metabolic syndrome, diabetes, or other lifestyle-related diseases. That is, LF-ES provides benefits to individuals who cannot perform adequate voluntary exercise as a consequence of excessive obesity and/or orthopedic problems. However, even though LF-ES is expected to be an alternative exercise method, it is also evident that blood pressure and cardiac afterload substantially rise during LF-ES. Our preliminary study in sedentary subjects revealed that systolic blood pressure (SBP) rose significantly, by 25 mm Hg on average, even at moderate LF-ES (3 METs). Furthermore, the magnitude of the elevation was quite different from subject to subject. For example, in one subject SBP was elevated by 50 mm Hg. The magnitude of SBP elevation might not matter for normal individuals, but it would cause * Corresponding author. Tel./fax: +81 75 753 6888. E-mail address: [email protected] (T. Moritani). 1050-6411/$ - see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.jelekin.2010.03.004

problems in some individuals with hypertension and/or myocardial infarction. Unfortunately, such symptoms seem to be common in prospective recipients of LF-ES, as mentioned above. Here, one of the major factors influencing SBP during LF-ES is the enhancement of intramuscular pressure in the contracted skeletal muscle. A large body of evidence has revealed that intramuscular pressure rises during muscle contraction, which results in the restriction of muscle blood flow (Aratow et al., 1993; Baumann et al., 1979; McDermott et al., 1982; Sejersted et al., 1984; Sjøgaard et al., 1986; Vedsted et al., 2006). That is, the elevation of intramuscular pressure by muscle contraction enhances peripheral vascular resistance (PVR), which creates resistance to the left ventricular ejection, resulting in the elevation of blood pressure. However, it is also reported that intramuscular pressure dose not remain elevated but fluctuates along with the cycle of skeletal muscle contraction/relaxation during intermittent muscle contractions (Aratow et al., 1993; Baumann et al., 1979; McDermott et al., 1982; Vedsted et al., 2006). That is, the PVR strictly rises and falls in synchronization with the cycle of muscle contraction and relaxation. In other words, the PVR rises only during the muscle contraction phase. These observations led us to speculate that the synchrony between cardiac systole and skeletal muscle contraction is an important determinant in blood pressure response during intermittent

T. Kimura et al. / Journal of Electromyography and Kinesiology 20 (2010) 572–579

muscle contractions. That is, if a cardiac systole and a skeletal muscle contraction became synchronized at a certain moment, the concomitant instantaneous SBP and cardiac afterload would rise as the enhanced PVR counteracted the left ventricular ejection. On the other hand, if a cardiac systole occurred asynchronously with a muscle contraction, the left ventricular ejection would occur during the lower PVR, lowering both the instantaneous SBP and the cardiac afterload. That is, from practical perspectives, it can be speculated that a well-designed LF-ES device, one that intentionally induces a muscle contraction asynchronously with a cardiac systole on each heartbeat, would allow lower SBP and reduced cardiac afterload during the entire duration of LF-ES. These assumptions have been suggested also as a physiological advantage of the ‘‘coupling of cardiac and locomotor rhythms” (Kirby et al., 1989b), which occurs during human locomotion (Kirby et al., 1989a,b; Niizeki, 2005; Niizeki et al., 1996). It was reported that the cardiac rhythm is synchronized with the locomotor cycle at a certain intensity during walking, running, and cycling, i.e., heartbeats are coupled with muscle contractions of each leg with a phase difference (Kirby et al., 1989a,b). These authors assumed that this coupling phenomenon might result in the maximization of blood flow to the exercising muscle and minimization of cardiac afterload if there is an appropriate phase difference between cardiac systole and muscle contraction. Niizeki et al. (1996) reported that when the subjects walked on a treadmill in synchrony with their ECG R-waves, SBP and diastolic blood pressure (DBP) seemed to be lowered. However, the attenuation of SBP and DBP was not clear and statistical analysis was not performed. In addition, the timing of each walking step was not precisely controlled, as the subject voluntarily walked in synchrony with the buzzer signal. That is, coupling’s effect on SBP and cardiac afterload has not been adequately examined, and therefore the details remain to be elucidated. Accordingly, the purposes of this study were (1) to investigate whether the phase difference between cardiac systole and skeletal muscle contraction (sc-s) affects instantaneous SBP and the concomitant cardiac afterload during intermittent muscle contractions; and (2) to determine, from practical perspectives, the most appropriate sc-s for the intermittent evoked muscle contraction (i.e., LF-ES) with lowest SBP and cardiac afterload. For these purposes, the sc-s was measured for every heartbeat during intermittent evoked muscle contractions and the concomitant instantaneous blood pressure, stroke volume (SV), PVR, and cardiac afterload were plotted as functions of sc-s. Our results would not only extend the knowledge about the physiological advantages of coordination between heart and skeletal muscle contractions, but also support the practical notion that a well-designed LF-ES device would allow lower SBP and cardiac afterload. 2. Methods 2.1. Subjects The present investigation was performed on eight healthy subjects (age 25.0 ± 2.8 yr, height 170.6 ± 2.0 cm, body mass 60.0 ± 2.7 kg, mean ± SE). The study protocol was approved by the Ethics Committee of Kyoto University and was performed in accordance with the Declaration of Helsinki. All subjects received an explanation of the nature and purpose of the study, and all gave their written informed consent to participate in it. The subjects were asked to avoid exercise, coffee, tea, and alcohol for 24 h before testing. 3. Experimental procedure Each subject, lying supine on a bed, received LF-ES for 15 min following a 5-min complete rest period. Rubber stimulation surface

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electrodes (5.0  7.5 cm) were applied over the quadriceps and biceps femoris muscles of each leg. A pair of electrodes was placed for each muscle group and the inter-electrode distance was secured for as long as possible to stimulate the whole motor units in the muscle group. The stimulator (SEN-7203, Nihon-Kohden, Tokyo, Japan) put out monophasic square waves (1.0 ms duration) through isolator devices (SS-104J, Nihon-Kohden). The stimulation frequency and duty cycle were 20 Hz and 0.3 s stimulation and 0.7 s pause, respectively. All muscle groups were synchronously stimulated so as to co-contract in an isometric manner, as this is a common method of LF-ES to prevent large movement of a lower limb. The stimulation intensity was individually set to the maximal level without pain in each subject (Hamada et al., 2003). The stimulation intensity was determined before the measurement. For each muscle group, the experimenter gradually increased the stimulation intensity until the subject reported pain. After the stimulation intensity was determined, a rest period was provided until the subject’s heart rate (HR) returned to the level before stimulation. The room temperature was kept at 25–26 °C. Each subject had visited our laboratory prior to the testing day in order to become familiar with the electrical muscle stimulation. 3.1. Data acquisition During the entire 20-min experiment, electrocardiography (ECG), beat-by-beat blood pressure (BP) and SV, and mechanomyogram (MMG) were continuously measured (Fig. 1). The ECG signal (CM5) was amplified and band-pass filtered between 1 and 100 Hz (BA-8224, Biotex, Kyoto, Japan). BP was continuously measured using arterial tonometry from the left hand (JENTOW-7700, Colin, Aichi, Japan). SV was recorded by impedance cardiography (AI601G, Nihon-Kohden, Tokyo, Japan). MMG, which detects minute dimensional changes in activated motor units, was recorded to monitor the cycle of muscle contraction and relaxation (Kimura et al., 2004; Orizio, 1993; Orizio et al., 1990; Shinohara and Søgaard, 2006; Yoshitake et al., 2002). The reason why we used an MMG device rather than conventional electromyography (EMG) is that the EMG cannot correctly record the action potentials of motor units because a large stimulation artifact saturates the EMG amplifier. Therefore, we employed the MMG, which is free of electrical artifacts and properly responds to muscle contraction. Our MMG recording method and procedures have been reported elsewhere (Kimura et al., 2004; Yoshitake et al., 2002). The MMG sensor (MP101-10, Medisens, Saitama, Japan) was composed of a small uniaxial accelerometer (base 9  9 mm, height 4.5 mm, weight 0.75 g, frequency response DC-1000 Hz). The MMG sensor was applied over the belly of the right rectus femoris muscle. The sensor was at the midpoint between the pair of stimulation electrodes. The sensor was secured to the skin with double-sided adhesive tape. The MMG signal was amplified and band-pass filtered between 1 and 250 Hz (MPS101, Medisens), differentiated, and full-wave rectified. The waveforms of ECG, BP, SV, and MMG were continuously digitized at 1 kHz (model 420, TransEra, Orem, UT, USA) and stored on a computer. In addition, gas exchange was measured by the mixing chamber method (AE-300S, Minato, Osaka, Japan) during the 20-min experiment. The VO2 was calculated on line every 15 s and stored on a computer. 3.2. Data analysis For the 5-min complete rest, average systolic and diastolic blood pressures (SBP and DBP) and HR between 1 and 4 min were calculated. Also, average SBP, DBP, and HR during 15-min muscle contractions were computed. In addition, for the 15-min muscle contractions, beat-by-beat SBP, DBP, cardiac afterload (tensiontime index, TTI), SV, and PVR were plotted as functions of the phase

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Subgroup 17 Subgroup 5 Subgroup 14 Subgroup 2 1sec Fig. 1. Example of the ECG, BP, impedance for SV measurement, and MMG, simultaneously recorded during intermittent electrical muscle contractions. The MMG signal was differentiated and full-wave rectified. In addition, example of the determination of the phase difference between cardiac systole and skeletal muscle contraction (sc-s) for each heartbeat was indicated.

difference between cardiac systole and skeletal muscle contraction (sc-s). As shown in Fig. 1, sc-s was defined as the time lag between the ECG R-peak and the onset of MMG. The positive value of sc-s means that the ECG R-peak occurred before the onset of MMG. On the contrary, the negative value of sc-s means that the ECG Rpeak occurred after the onset of MMG. And sc-s was always greater than or equal to 500 ms and less than 500 ms, as the length of the stimulation duty cycle was 1000 ms (300 ms stimulation and 700 ms pause). For example, R-peak(53) (i.e., 53rd R-peak) in Fig. 1 occurred 332 ms prior to the onset of MMG, hence sc-s of this R-peak (i.e., sc-s(53)) was 332 ms. On the other hand, sc-s(54) was 275 ms because R-peak(54) occurred 275 ms after the onset of MMG. Thus, sc-s was calculated for all heartbeats during 15-min intermittent muscle contractions. Then, depending on the length of sc-s, all heartbeats were categorized into 20 subgroups, which were sectioned by 50-ms intervals between 500 ms and 500 ms. That is, subgroup-a (a = 1, 2, 3 . . . 20) was composed of heartbeats whose sc-s were greater than or equal to ( 500 + 50  (a 1)) ms and less than ( 450 + 50  (a 1)) ms. For example, sc-s(54) in Fig. 1 was 275 ms, hence this heartbeat (i.e., Rpeak(54)) was categorized into subgroup 5 ( 300 ms <= sc-s < 250 ms). Thus, all heartbeats were categorized into 20 subgroups. In addition, as shown in Fig. 2, when each heartbeat was categorized into one of 20 subgroups, the BP and SV signals associated

with the heartbeat were also categorized into the same subgroup. And the BP and SV signals were respectively trigger-averaged using the ECG R-peak with BP and SV signals at the other heartbeats in the same subgroup. The lengths of the sampling windows for BP and SV signals were 1 s and the starting points of the windows corresponded to the occurrence of the R-wave. For example, in Fig. 2, the raw signals of BP and SV associated with R-peak(54) were, respectively, categorized into subgroup 5 because the R-peak(54) had been categorized into subgroup 5. Then, the BP and SV signals were respectively trigger-averaged with the other BP and SV signals in subgroup 5. Thus, after all of the raw signals were trigger-averaged in the respective subgroups, the SBP, DBP, and SV were computed from the averaged signals in each subgroup. In addition, the tension-time index (TTI) was calculated from each averaged BP waveform to evaluate the instantaneous cardiac afterload (Barnard et al., 1973, 1977; Tarazi and Levy, 1982). In order to obtain TTI, area under the BP waveform was computed for ejection phase and the area was multiplied by the average heart rate. Also, the instantaneous PVR was estimated as the ratio of mean blood pressure (MBP) to SV in each subgroup. The MBP was calculated as DBP + (SBP-DBP)/3. Fig. 3 shows a typical example of SBP response as a function of sc-s from a single subject. Thus, SBP, DBP, TTI, SV, and PVR were respectively calculated for each of 20 subgroups.

T. Kimura et al. / Journal of Electromyography and Kinesiology 20 (2010) 572–579

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R-peak(54)

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τc-s (54) = -275ms Subgroup 5 1sec Fig. 2. Example of the trigger averaging of BP and SV waveforms after categorization into the subgroup. Note that the lengths of sampling windows for BV and SV are changed for visual purposes. See text for details.

nificance was set at P < 0.05. Descriptive statistics included mean and SE.

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4. Results

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τ c-s c s (ms) Fig. 3. Example of the instantaneous SBP response as a function of sc-s from a subject. His ECG waveform is overwritten as a reference. The negative value of sc-s means that the skeletal muscle contraction began before the ECG R-peak. The 500 ms in the X-axis indicates the subgroup of 500 ms <= sc-s < 450 ms.

3.3. Statistical analysis A one-way ANOVA with repeated-measures and Tukey’s post hoc test was used to assess whether or not sc-s significantly changed SBP, DBP, TTI, SV, and PVR. Student’s paired t-test was employed to determine whether or not VO2 and HR were significantly changed by evoked muscle contraction. Statistical sig-

The average VO2 across subjects was 3.89 ± 0.17 ml/min/kg and 7.88 ± 0.75 ml/min/kg during rest and evoked muscle contraction, respectively. And the average HR was 60.5 ± 2.7 beats/min and 78.7 ± 5.1 beats/min during rest and evoked muscle contraction, respectively. The VO2 and HR were significantly increased by muscle contraction (P < 0.05, paired t-test). The average number of heartbeats in each subgroup was 58.4 ± 1.0 on average across subjects. In one subject, the TTI could not be determined because of his distinct BP waveform. Therefore, TTI was evaluated in seven subjects. Fig. 4 represents the responses of instantaneous SBP, TTI, and SV as functions of sc-s, averaged across subjects. It shows that SBP and TTI rise at the negative sc-s, i.e., when muscle contraction starts before the ECG R-peak, and falls at the positive sc-s, i.e., when muscle contraction occurs after the R-peak. One-way ANOVA indicated that each of two parameters was significantly changed by sc-s (P < 0.05). And Tukey’s post hoc test confirmed that there was a significant difference between the values at 0–50 ms and those at negative sc-s in each parameter (P < 0.05). Very similar results were obtained in DBP and PVR (not shown in figure). The response of instantaneous SV averaged across subjects, is also indicated in Fig. 4. It was found that SV was lower at the smaller sc-s (from 450 to 250 ms) and enhanced at positive sc-s. One-way ANOVA demonstrated that SV was significantly changed by sc-s (P < 0.05), and Tukey’s post hoc test revealed a significant difference between the values at 0–50 ms and those at negative sc-s (P < 0.05). Fig. 5 represents average SBP and DBP of each subject at the 5min complete rest and at the 15-min muscle contractions (LF-ES). It represents that the magnitude of the elevation of SBP was substantially different from subject to subject. Especially, in two subjects SBP rose to over 160 mm Hg. Therefore, we divided the subjects into two groups based on the SBP at LF-ES, i.e., low SBP group (n = 4) and high SBP group (n = 4), as indicated in Fig. 5.

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Fig. 4. The responses of instantaneous SBP, TTI, and SV as functions of sc-s, averaged across subjects. The 500 ms in the X-axis indicates the subgroup of sc-s < 450 ms. *P < 0.05, significantly different from the value at 0–50 ms. Results are expressed as mean and SE.

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Fig. 5. The SBP and DBP of each subject at the 5-min complete rest (Rest) and at the 15-min muscle contractions (LF-ES). The subjects were divided into two groups, i.e., low SBP group (n = 4, triangle symbol) and high SBP group (n = 4, square symbol).

Then, the responses of instantaneous SBP, DBP, and PVR as functions of sc-s were averaged in each group (Fig. 6). The results showed that the amplitudes of SBP and DBP fluctuations were much greater in high SBP group. Very similar result was obtained in TTI (not shown in figure). In addition, as shown in Fig. 6, PVR of high SBP group highly increased at negative sc-s. Thus, the impact of sc-s on hemodynamic responses was much greater in high SBP group.

5. Discussion This study’s novel finding is that the instantaneous blood pressure and cardiac afterload depend strongly on the phase difference between cardiac systole and muscle contraction (sc-s) during intermittent muscle contractions. The results show that when a muscle contraction starts after an ECG R-peak, the SBP and TTI are significantly lowered. Our hypothesis was that if a muscle contraction

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Fig. 6. The responses of instantaneous SBP, DBP, and PVR as a function of sc-s, averaged in low SBP group (dashed line) and high SBP group (solid line). The indicates the subgroup of 500 ms <= sc-s < 450 ms.

occurred asynchronously with a cardiac systole, left ventricular ejection would occur during the skeletal muscle relaxation phase, thus lowering both instantaneous SBP and the cardiac afterload. Therefore, the present results support our hypothesis. However, strictly speaking, the present results also indicate that when a muscle contraction starts even just after an ECG R-peak (i.e., 0 ms <= sc-s < 50 ms), SBP and TTI are significantly lower. The ECG QRS complex indicates ventricular depolarization followed by left ventricular ejection. Therefore, it might be argued that the present result does not support our hypothesis; rather, it demonstrates that the fully simultaneous contractions of cardiac and skeletal muscles result in lower SBP and cardiac afterload. However, it should be taken into account that there is a considerable time lag between the initiation of skeletal muscle contraction and the occurrence of peak intramuscular pressure. It has been reported that the time to peak force in maximal twitch contraction of quadriceps muscles is 73.5 ms (Hamada et al., 2000) or 87.5 ms (Verges et al., 2009). Furthermore, it should be noted that a repetitive stimulation (three stimulation pulses at 10 Hz), which is similar to the stimulation pattern in the present study, resulted in a much longer time to peak force (168.7 ms) in quadriceps muscles (Verges et al., 2009). The time to peak force of biceps femoris muscles has never been reported. However, it has been elucidated that the proportion of slow type I fibers of biceps femoris muscles was substantially higher than that of quadriceps muscles (Johnson et al., 1973). Therefore, it can be speculated that the time to peak force of biceps femoris would be at least as long as that of quadriceps. On the other hand, we have revealed that the MMG is very sensitive and can detect minute dimensional changes in muscle fibers induced by a twitch contraction of a single motor unit (Yoshitake et al., 2002). Therefore, at the moment at which the MMG onset was detected, a few motor units, located very near the skin surface, started to contract at the very beginning. Therefore, it is reasonable to think that a substantial time lag existed between the onset of MMG and the peak of the muscle contraction force in each contraction phase. This indicates that the highest intramuscular pressure would have occurred substantially after the onset of MMG in each contraction phase. This could be a reason why SBP and cardiac afterload were significantly lower when the sc-s was between 0

500 ms in X-axis

and 50 ms. Therefore, the present result properly supports our hypothesis that the asynchronicity between a cardiac systole and a rise in intramuscular pressure results in lower concomitant SBP and cardiac afterload. On the other hand, direct measurement of intramuscular pressure would have verified our hypothesis more clearly. However, measurement of intramuscular pressure is always invasive (e.g., catheter technique) (Aratow et al., 1993; McDermott et al., 1982; Sjøgaard et al., 1986; Vedsted et al., 2006). Therefore, we did not measure intramuscular pressure because it causes pain, which largely influences the blood pressure response. In addition, as mentioned in the introduction, our final goal was to determine the most appropriate phase difference for a well-designed stimulator, one that induces muscle contraction coupling with the heartbeat. Therefore, assessment of appropriate phase difference between the ECG R-peak and the onset of muscle contraction is sufficient to achieve our goal. Previous studies reported that the cardiac rhythm is synchronized with the locomotor cycle at a certain intensity during walking, running, cycling, and rhythmic cuff occlusion of the thigh (Kirby et al., 1989a,b; Niizeki, 2005). The authors assumed that this coupling phenomenon contributes to the maximization of blood flow to the muscle and thus minimizes cardiac afterload if there is an appropriate phase difference between cardiac systole and muscle contraction. However, the effects of the phase difference on cardiac afterload and SBP have not been adequately examined. Only one study has addressed this assumption, in which three subjects walked on a treadmill in synchrony with their ECG R-waves with four fixed phase delays (Niizeki et al., 1996). Those authors suggested that, similar to our present results, SBP and DBP seemed to be lowered when the subjects walked with zero phase delay (i.e., when they were completely synchronized with ECG R-wave). However, the attenuation of SBP and DBP was not clear and statistical analysis was not performed. In addition, the phase difference was not precisely controlled, as the subject voluntarily walked in synchrony with the buzzer signal. On the other hand, we statistically demonstrated that the phase difference (sc-s) significantly affects the SBP and DBP, with much higher time resolution. Furthermore, we found for the first time that phase difference significantly affected cardiac afterload, stroke volume, and PVR.

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In terms of clinical relevance, the present results support the utility of a well-designed stimulator, one that induces the coupling of muscle contraction with the heartbeat with an appropriate phase difference (i.e., positive sc-s), to allow lower SBP and reduce cardiac afterload during the entire duration of LF-ES. LF-ES has been shown to be an alternative exercise prescription for individuals who cannot perform adequate voluntary exercise as a consequence of excessive obesity and/or orthopedic problems (Banerjee et al., 2005; Caulfield et al., 2004; Hamada et al., 2003, 2004). Therefore, it can be assumed that a lot of prospective recipients of LF-ES suffer from hypertension and/or myocardial infarction. In that regard, the present result would be an important finding for clinical application. Furthermore, the effects of the phase difference (sc-s) on blood pressure and cardiac afterload were much greater in high SBP group (Fig. 6). This further supports that the new LF-ES device would be more effective in individuals with higher blood pressure. Fig. 7 illustrates an example of a specially designed LF-ES in practical use, based on the results of the present study. That is, the recipient’s ECG is continuously monitored on line and the ECG QRS complex in each heartbeat is used as a trigger of an output to the stimulator. The stimulator generates a train of stimulation pulses, each of which occurs when the stimulator receives an impulse from the ECG monitor. Thus, this LF-ES device induces muscle contraction coupling with the heartbeat with an appropriate phase difference (i.e., positive sc-s), resulting in lower SBP and reduced cardiac afterload during the entire duration of LF-ES. However, there is a concern that the HR highly increases in some patients, even though the increase of HR was moderate (about 20 beats/min) in the healthy subject of the present study. In such a patient, it might be difficult to apply this stimulation method because the time interval between muscle contractions becomes very short. Further investigation is needed to resolve this problem. For example, it might be effective if the right and left legs are alternately contracted as voluntary running or cycling. Then, the heartbeat can be coupled with the muscle contraction of both legs as seen during voluntary running and cycling (Kirby et al., 1989a,b), i.e., if the subject’s HR is 120 beats/min, each leg will be contracted at the rate of 60 times/min. In summary, the present study demonstrated that the phase difference between cardiac systole and muscle contraction strongly affects concomitant blood pressure and cardiac afterload during intermittent muscle contractions. This result sup-

Trigger on

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Stimulator

Fig. 7. Schematic diagram of the LF-ES triggered by the ECG R-waves. See text for details.

ports the notion that a well-designed stimulator, one that induces muscle contraction coupling with the heartbeat with an appropriate phase difference, would allow lower SBP and cardiac afterload during the entire duration of LF-ES for clinical applications. Acknowledgments This study was supported by a Grant-in-Aid for Scientific Research from the Japan Society for the Promotion of Science (No. 18300230). References Aratow M, Ballard RE, Crenshaw AG, Styf J, Watenpaugh DE, Kahan NJ, et al. Intramuscular pressure and electromyography as indexes of force during isokinetic exercise. J Appl Physiol 1993;74(6):2634–40. Banerjee P, Clark A, Witte K, Crowe L, Caulfield B. Electrical stimulation of unloaded muscles causes cardiovascular exercise by increasing oxygen demand. Eur J Cardiovasc Prev Rehabil 2005;12(5):503–8. Barnard RJ, MacAlpin R, Kattus AA, Buckberg GD. Ischemic response to sudden strenuous exercise in healthy men. Circulation 1973;48(5):936–42. Barnard RJ, MacAlpin R, Kattus AA, Buckberg GD. Effect on training on myocardial oxygen supply/demand balance. Circulation 1977;56(2):289–91. Baumann JU, Sutherland DH, Hänggi A. Intramuscular pressure during walking: an experimental study using the wick catheter technique. Clin Orthop Relat Res 1979;145:292–9. Caulfield B, Crowe L, Minogue C, Banerjee P, Clark A. The use of electrical muscle stimulation to elicit a cardiovascular exercise response without joint loading: a case study. J Exerc Physiol Online 2004;7(3):84–8. Hamada T, Hayashi T, Kimura T, Nakao K, Moritani T. Electrical stimulation of human lower extremities enhances energy consumption, carbohydrate oxidation, and whole body glucose uptake. J Appl Physiol 2004;96(3):911–6. Hamada T, Sale DG, MacDougall D, Tranopolsky MA. Postactivation potentiation, fiber type, and twitch contraction time in human knee extensor muscles. J Appl Physiol 2000;88(6):2131–7. Hamada T, Sasaki H, Hayashi T, Moritani T, Nakao K. Enhancement of whole body glucose uptake during and after human skeletal muscle low-frequency electrical stimulation. J Appl Physiol 2003;94(6):2107–12. Johnson MA, Polgar J, Weightman D, Appleton D. Data on the distribution of fibre types in 36 human muscles. J Neurol Sci 1973;18(1):111–29. Kimura T, Hamada T, Watanabe T, Maeda A, Oya T, Moritani T. Mechanomyographic responses in human biceps brachii and soleus during sustained isometric contraction. Eur J Appl Physiol 2004;92(4–5):533–9. Kirby RL, MacLeod DA, Marble AE. Coupling between cardiac and locomotor rhythms: the phase lag between heart beats and pedal thrusts. Angiology 1989a;40(7):620–5. Kirby RL, Nugent ST, Marlow RW, MacLeod DA, Marble AE. Coupling of cardiac and locomotor rhythms. J Appl Physiol 1989b;66(1):323–9. McDermott AG, Marble AE, Yabsley RH, Phillips MB. Monitoring dynamic anterior compartment pressures during exercise. A new technique using the STIC catheter. Am J Sports Med 1982;10(2):83–9. Niizeki K. Intramuscular pressure-induced inhibition of cardiac contraction: implications for cardiac-locomotor synchronization. Am J Physiol Regul Integr Comp Physiol 2005;288(3):R645–50. Niizeki K, Kawahara K, Miyamoto Y. Cardiac, respiratory, and locomotor coordination during walking in humans. Folia Primatol (Basel) 1996;66(1– 4):226–39. Orizio C. Muscle sound: bases for the introduction of a mechanomyographic signal in muscle studies. Crit Rev Biomed Eng 1993;21(3):201–43. Orizio C, Perini R, Diemont B, Maranzana Figini M, Veicsteinas A. Spectral analysis of muscular sound during isometric contraction of biceps brachii. J Appl Physiol 1990;68(2):508–12. Sejersted OM, Hargens AR, Kardel KR, Blom P, Jensen O, Hermansen L. Intramuscular fluid pressure during isometric contraction of human skeletal muscle. J Appl Physiol 1984;56(2):287–95. Shinohara M, Søgaard K. Mechanomyography for studying force fluctuations and muscle fatigue. Exerc Sport Sci Rev 2006;34(2):59–64. Sjøgaard G, Kiens B, Jørgensen K, Saltin B. Intramuscular pressure, EMG and blood flow during low-level prolonged static contraction in man. Acta Physiol Scand 1986;128(3):475–84. Tarazi RC, Levy MN. Cardiac responses to increased afterload. Hypertension 1982;4(Suppl. 2):8–18. Vedsted P, Blangsted AK, Søgaard K, Orizio C, Sjøgaard G. Muscle tissue oxygenation, pressure, electrical, and mechanical responses during dynamic and static voluntary contractions. Eur J Appl Physiol 2006;96(2):165–77. Verges S, Maffiuletti NA, Kerherve H, Decorte N, Wuyam B, Millet GY. Comparison of electrical and magnetic stimulations to assess quadriceps muscle function. J Appl Physiol 2009;106(2):701–10. Yoshitake Y, Shinohara M, Ue H, Moritani T. Characteristics of surface mechanomyogram are dependent on development of fusion of motor units in humans. J Appl Physiol 2002;93(5):1744–52.

T. Kimura et al. / Journal of Electromyography and Kinesiology 20 (2010) 572–579 Tetsuya Kimura received his Ph.D. degree in Human and Environmental Studies from Kyoto University in 2007. Currently he is a post doctoral fellow at Kyoto University under the direction of Dr. Moritani. His current research interest includes the analysis of activation strategies of motor units during human fundamental motion and electrical muscle stimulation for enhancing energy expenditure.

Naoko Kameda is a Research Associate of Department of Food and Nutrition Science at Sagami Women’s Junior College. She graduated at the Department of Nutritional Science, Okayama Prefectural University in 2007. She worked in the Laboratory of Applied Physiology of the Graduate School of Human and Environmental Studies, Kyoto University under the direction of Dr. Toshio Moritani and received M.Sc. degree in 2009. She is currently investigating the effect of nutritional elements ingestion on autonomic nervous system.

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Toshio Moritani was born in Japan in 1950. He received his Ph.D. degree in Sports Medicine from the University of Southern California in 1980 under the direction of Dr. Herbert A. deVries. In 1985, following faculty appointments at the University of Texas at Arlington and Texas A&M University, he returned to Japan and joined the Department of Integrated Human Studies at Kyoto University. In 1992, he was appointed Associate Professor of Applied Physiology at the Graduate School of Human and Environmental Studies at Kyoto University and became Professor since 2000. He is currently Director of the Laboratory of Applied Physiology.