Cadence-dependent changes in corticospinal excitability of the biceps brachii during arm cycling

Abstract

This is the first study to report the influence of different cadences on the modulation of supraspinal and spinal excitability during arm cycling. Supraspinal and spinal excitability were assessed using transcranial magnetic stimulation of the motor cortex and transmastoid electrical stimulation of the corticospinal tract, respectively. Transcranial magnetic stimulation-induced motor evoked potentials and transmastoid electrical stimulation-induced cervicomedullary evoked potentials (CMEPs) were recorded from the biceps brachii at two separate positions corresponding to elbow flexion and extension (6 and 12 o'clock relative to a clock face, respectively) while arm cycling at 30, 60 and 90 rpm. Motor evoked potential amplitudes increased significantly as cadence increased during both elbow flexion (P < 0.001) and extension (P = 0.027). CMEP amplitudes also increased with cadence during elbow flexion (P < 0.01); however, the opposite occurred during elbow extension (i.e., decreased CMEP amplitude; P = 0.01). The data indicate an overall increase in the excitability of corticospinal neurons which ultimately project to biceps brachii throughout arm cycling as cadence increased. Conversely, changes in spinal excitability as cadence increased were phase dependent (i.e., increased during elbow flexion and decreased during elbow extension). Phase- and cadence-dependent changes in spinal excitability are suggested to be mediated via changes in the balance of excitatory and inhibitory synaptic input to the motor pool, as opposed to changes in the intrinsic properties of spinal motoneurons.

in animals, complex well-coordinated locomotor activities (e.g., fictive locomotion and scratch in the adult decerebrate cat) can be evoked in the mammalian spinal cord without the influences of supraspinal input and/or sensory feedback. The basic alternating rhythm of these motor outputs is generated by a network of cells located within the spinal cord, referred to as the central pattern generator (CPG) (Grillner 1981; Jordan 1998). In the adult decerebrate cat and rat, the intrinsic properties of spinal motoneurons are altered during CPG-mediated motor outputs (i.e., fictive scratch and locomotion), such that they become more excitable prior to and throughout the motor task (Brownstone et al. 1992, 1994; Krawitz et al. 2001; MacDonell et al. 2015; Power et al. 2010). It is currently unclear whether increases in spinal motoneuron excitability during CPG-mediated motor outputs are influenced by cadence.

In humans, indirect evidence suggests that, similar to quadrupeds, rhythmic and alternating locomotor outputs such as cycling (leg and arm) are also mediated, in part, by a spinal CPG (Capaday et al. 1999; Carroll et al. 2006; Pyndt and Nielsen 2003; Zehr et al. 2009; Zehr and Stein 1999), although supraspinal input to the spinal cord is also required (Forman et al. 2014; Petersen et al. 2001; Sidhu et al. 2012). The majority of studies involving the assessment of nervous system excitability during locomotor outputs focus on the modulation of supraspinal and/or spinal reflex excitability, using a set cadence. Moreover, those studies that involve changes in cadence (e.g., walking vs. running) tend to focus on the processing of sensory information. For example, suppression of soleus H-reflex gain has been reported with running compared with walking (Capaday and Stein 1987; Ferris et al. 2001; Simonsen and Dyhre-Poulsen 1999). With the use of cycling as a model of locomotor output, the amplitudes of somatosensory evoked potentials and H-reflexes are suppressed as leg cycling cadence increases (Staines et al. 1997). Decreases in reciprocal inhibition between antagonistic muscles also occur as leg cycling cadence increases (Pyndt et al. 2003). While it is evident that cadence or velocity can alter the processing of sensory information, it is unclear what effect this has on the output of the motor system.

One way to examine the effect(s) of locomotor cadence or velocity on motor system output is to assess the excitability of the corticospinal pathway. Composed of the motor cortex and spinal motoneurons, the corticospinal pathway is a major descending motor pathway involved in the voluntary control of human motor output. A common means to examine corticospinal excitability (CSE) is to assess the amplitude of motor evoked potentials (MEPs) elicited via transcranial magnetic stimulation (TMS) of the motor cortex. TMS-evoked MEPs provide information regarding the excitability of the whole corticospinal pathway and can thus be influenced by changes in supraspinal and/or spinal excitability. Transmastoid electrical stimulation (TMES) of corticospinal axons, which elicits a cervicomedullary-MEP (CMEP), can be used to assess spinal excitability, allowing one to make inferences as to where along the corticospinal pathway (supraspinal or spinal) changes in CSE may have occurred. During arm cycling at a set cadence (60 rpm) and workload (25 W), our laboratory recently reported that CSE of the biceps brachii was higher during arm cycling than an intensity-matched tonic contraction with spinal excitability of the biceps brachii being higher at the initiation of flexion (Forman et al. 2014). In that same study, we also probed the effects of increased cycling cadence on spinal excitability of the biceps brachii during the extension phase of arm cycling when the elbow was extending and the triceps brachii became the prime mover (Forman et al. 2014). There was a slight, albeit nonsignificant (likely due to the low sample size of 4), decrease in CMEP amplitude at 90 rpm compared with 60 rpm. In fact, spinal excitability of the biceps brachii was reduced during the extension phase of cycling at both 60 and 90 rpm compared with a position-matched rest condition (i.e., elbow extended) (Forman et al. 2014). To the best of our knowledge, human studies have not examined the effects of altered cycling cadence on transient, phase-dependent changes in CSE (i.e., supraspinal and spinal) during locomotor outputs such as arm cycling.

The purpose of the present study was to examine the influence of cadence on CSE of the biceps brachii during the elbow flexion and extension phases of arm cycling. We assessed corticospinal and spinal excitability via TMS-evoked MEPs and TMES-evoked CMEPs recorded from the biceps brachii, by altering the cadence of arm cycling at a constant workload. We hypothesized that increased cycling cadence would increase CSE throughout arm cycling (i.e., during both flexion and extension) due in large part to increases in spinal excitability.

METHODS

Ethical Approval

The procedures of the experiment were verbally explained to each volunteer prior to the start of the session. Once all questions were answered, written consent was obtained. This study was conducted in accordance with the Helsinki declaration and approved by the Interdisciplinary Committee on Ethics in Human Research at Memorial University of Newfoundland (ICEHR no. 201413110-HK). Procedures were in accordance with the Tri-Council guideline in Canada, and potential risks were fully disclosed to participants.

Participants

Ten male volunteers (24.3 ± 5.4 yr of age, 177.8 ± 5.73 cm, 81.7 ± 7.3 kg, eight right-hand dominant, two left-hand dominant) partook in this study. Participants had no known neurological impairments. Prior to the experiment, all volunteers completed a magnetic stimulation safety checklist to screen for contraindications to magnetic stimulation (Rossi et al. 2009). Additionally, participants were required to complete a Physical Activity Readiness Questionnaire to screen for any contraindications to exercise or physical activity.

Experimental Setup

This study was carried out on an arm cycle ergometer (SCIFIT ergometer, model PRO2 Total Body). Participants were seated upright at a comfortable distance from the hand pedals, so that, during cycling, there was no reaching or variation in trunk posture. To further ensure that posture was maintained throughout all trials, each participant was strapped securely to the ergometer seat with straps placed over the shoulders and criss-crossed over the chest. Movement of the shoulders and arms was not impeded. The hand pedals of the ergometer were fixed 180° out of phase, and the seat height was adjusted so that the shoulders of each individual were approximately the same height as the center of arm crank shaft. Participants lightly gripped the ergometer handles with the forearms pronated and wore wrist braces to limit the movement of the wrists during cycling as heteronymous reflex connections exist between the wrist flexors and biceps brachii (Manning and Bawa 2011).

Measurements were taken from two different locations: 6 and 12 o'clock relative to a clock face, whereby 12 o'clock was defined as the “top dead center” of the arm crank, and 6 o'clock was defined as the “bottom dead center.” These sites were relative to the hand dominance of each individual. For example, 12 o'clock for a right-handed participant would have been when their right hand was positioned at “top dead center” of the arm crank. For a left-handed individual, 12 o'clock would have been set when their left hand was at “top dead center.” These two positions were chosen as they represent periods of high (6 o'clock) and low (12 o'clock) levels of biceps brachii (the main muscle of interest) activation during arm cycling (demonstrated in Fig. 1). Movement between 3 o'clock (when the elbow reaches full extension) and 9 o'clock (when the elbow reaches maximal flexion) occurs when the elbow is flexing and the biceps brachii is most active. Movement between 9 o'clock and 3 o'clock occurs when the elbow is extending and the biceps brachii is less active. Measurements at each position were taken separately.

Fig. 1.

A: rectified electromyographic (EMG) values of the biceps brachii and triceps brachii from a single participant throughout one full revolution of arm cycling. The 1 denotes the 6 o'clock position, where the elbow is flexing and the biceps brachii is active, and 2 occurs during elbow extension, when the biceps brachii is relatively inactive. Averaged EMG activity of the biceps brachii (B) and triceps brachii (C) throughout arm cycling at varying cadences is shown. The dashed, light gray lines represent cycling at 30 rpm; dark, solid gray lines, cycling at 60 rpm; and black solid lines are cycling at 90 rpm. Each crank position represents 8.3% of a full cycling revolution. EMG rectified averages were normalized to the largest EMG elicited between the three different cadences.

The study required participants to cycle at three different cycling cadences, 30, 60 and 90 rpm, at a constant workload of 20 W. Measurements were taken at 6 and 12 o'clock for a total of six separate trials. The order of these trials was randomized, and responses (described below) were triggered automatically when the arm crank passed by one of the two predetermined positions.

Electromyography Recordings

Electromyographic (EMG) activity of the biceps brachii and triceps brachii of the dominant arm were recorded using pairs of surface electrodes (Medi-Trace 130 ECG conductive adhesive electrodes) positioned over the midline of the biceps brachii and the lateral head of the triceps brachii. A ground electrode was placed on the lateral epicondyle. Prior to electrode placement, the skin was thoroughly prepared by removal of dead epithelial cells (using abrasive paper), followed by sanitization with an isopropyl alcohol swab. EMG was collected on-line at 5 kHz using CED 1401 interface and Signal 4 [Cambridge Electronic Design (CED), Cambridge, UK] software program. Signals were amplified (gain of 300) and filtered using a three-pole Butterworth with cutoff frequencies of 10-1,000 Hz.

Stimulation Conditions

Motor responses from the biceps brachii were elicited via 1) electrical stimulation at Erb's point; 2) TMS; and 3) TMES. All volunteers had prior experience with TMS, TMES and Erb's point stimulation procedures. While determining stimulation intensities, participants were instructed to place their hands lightly on the hand pedals of the ergometer. Their dominant hand was placed at the 6 o'clock position, and their nondominant hand at the 12 o'clock position. For the purpose of this study, this was considered “rest”.

Brachial Plexus Stimulation

The maximal M-wave (M_max) of the biceps brachii was first determined by eliciting M-waves through electrical stimulation of the brachial plexus at Erb's point (DS7AH, Digitimer, Welwyn Garden City, Hertfordshire, UK). A pulse duration of 200 μs was used, and intensities ranged from 100–300 mA. The cathode was placed in the supraclavicular fossa, and the anode on the acromion process. The initial stimulation intensity was set at 25 mA and gradually increased until the elicited M-waves of the biceps brachii reached a plateau. Stimulation intensity was then increased by 10% to ensure M_max were elicited throughout the study. Following analysis, MEP and CMEP amplitudes were normalized to the M_max during each trial to account for changes in peripheral neuromuscular propagation (Taylor 2006).

TMS

MEPs were elicited via TMS with the use of a Magstim 200 (Magstim, Dyfed, UK). Stimulations were delivered over the vertex via a circular coil (13.5-cm outside diameter). Vertex was determined by measuring the midpoint between the participant's nasion and inion, and the midpoint between the participant's tragi. The intersection of these two points was measured, marked and defined as vertex (Copithorne et al. 2015; Forman et al. 2014; Pearcey et al. 2014; Philpott et al. 2015; Power and Copithorne 2013). The coil was held tangentially to the participant's skull, approximately parallel to the floor, with the direction of the current flow preferentially activating either the left or right motor cortex (depending on hand dominance). The coil was held firmly against the participant's head by one of the investigators to ensure careful and consistent alignment over vertex for each trial. Stimulation intensity was started at ∼25% of magnetic stimulator output (MSO) and gradually increased until motor threshold was found. Motor threshold was defined as the lowest %MSO that resulted in a MEP amplitude of 50 μV or greater in 50% of trials (4 out of 8). This %MSO was then increased by 20%, and an average of eight MEPs, elicited at “rest” at this new intensity, was then calculated. This %MSO was used throughout the remainder of the experiment.

TMES

TMES was delivered using Ag-AgCl surface electrodes applied just inferior to the mastoid processes. The pulse duration was fixed at 100 μs, and stimulations intensities of 125–275 mA were used (DS7AH, Digitimer, Welwyn Garden City, Hertfordshire, UK). Stimulation intensity began at 25 mA and gradually increased until the average of eight CMEP amplitudes matched the average of the eight MEP amplitudes previously determined. This stimulation intensity was used throughout the remainder of the experiment.

Experimental Protocol

Once the intensities for Erb's point stimulation, TMS, and TMES were determined, the six different cadence trials (30, 60 and 90 rpm at 6 and 12 o'clock) were performed. The arm ergometer was set to a fixed workload of 20 W, and participants were instructed to cycle at a specified cadence. While cycling, a trial consisting of 10 MEPs and 10 CMEPs was completed at one of two positions (6 or 12 o'clock). The order of these stimulations was randomized throughout the trial, and stimulations were separated by ∼7–8 s. The total length of cycling was 2.5 min. To account for possible changes in the compound muscle action potential, a second trial consisting of three M-waves was performed immediately thereafter, given that M_max may change over the course of an experiment (Crone et al. 1999). These stimulations were elicited at the same workload, cadence and position of the previous MEPs and CMEPs. They were also separated by 7–8 s. These steps were then repeated for the remaining five trials.

Measurements

Data were analyzed off-line using Signal 4 software (CED). The peak-to-peak amplitudes of MEPs, CMEPs and M_max of the biceps brachii were measured. The peak-to-peak amplitudes for all evoked potentials were measured from the initial deflection of the voltage trace from the baseline EMG to the return of the trace to baseline levels. Because changes in MEP and CMEP amplitudes can be the result of changes to M_max, both MEPs and CMEPs were normalized to the M_max evoked during the same trial. Prestimulus EMG, defined as a window of the mean rectified EMG immediately prior to the stimulation artifact, was measured from the rectified traces. The length of the window was determined by the cadence of the trial: 30 rpm, 100 ms; 60 rpm, 50 ms; 90 rpm, 33.3 ms. Thus each timeframe from which the background EMG (bEMG) measurement was made represented 5% of the total time it took to complete one revolution. Measurements were taken from the averaged files of all 10 CMEPs, 10 MEPs and 5 M_max.

Statistics

All statistical analysis was performed using IBM's SPSS Statistics version 19. Separate one-way (i.e., cadence) repeated-measures ANOVAs were used to assess whether statistically significant differences in MEP or CMEP amplitudes (normalized to M_max) and the average of the prestimulus EMG occurred between the three cycling cadences at each phase of the cycle (i.e., elbow flexion and extension). All statistics were run on group data, and a significance level of P < 0.05 was used. All data are reported in text as means ± SD and illustrated in figures as means ± SE. Although we collected all data during different cycling phases within each cadence, we did not use a two-way repeated-measures ANOVA (phase × cadence) for CSE, given that our laboratory has previously shown phase-dependent differences in CSE (Forman et al. 2014). Differences between these positions were expected, given that they represent the active propulsion phase (elbow flexion) and the recovery phase (elbow extension) of the biceps brachii. Our interest was in assessing changes in CSE within a cycling phase as a function of cadence.

To make inferences as to changes in supraspinal and spinal excitability during cycling, it is important that the intensity of the motor output, as estimated via bEMG levels, in the biceps brachii and triceps brachii be similar when MEPs and CMEPs are evoked. Thus we compared bEMG levels using a two-way repeated-measures ANOVA with factors of “stimulation type” and “cadence” (2 × 3).

RESULTS

M_max Recorded from Biceps Brachii as Function of Cycling Cadence and Elbow Position

There was no significant main effect for “cadence” (P = 0.095) or “elbow position” (P = 0.109), nor was there an interaction of “cadence” and “elbow position” (P = 0.50) on M_max recorded from the biceps brachii.

EMG Patterns during Arm Cycling

The rectified EMG of a single participant for the biceps brachii and triceps brachii during arm cycling is shown in Fig. 1. In this representative example, 10 frames without stimulation were rectified and averaged over 1 s, with the participant cycling at a cadence of 60 rpm. At this cadence, 1 s represents one full revolution of arm cycling. The black arrows labeled 1 and 2 denote the two points where stimulations would have occurred had they been delivered (flexion and extension, respectively). Average values of cycling EMG from all participants during arm cycling are plotted for the biceps and triceps brachii in Figs. 1, B and C, respectively. This was done for all 12 positions, which were made relative to a clock face.

Cadence-dependent Changes in CSE during the Flexion Phase of Arm Cycling

CSE during elbow flexion.

MEP AMPLITUDE.

Figure 2A shows the average of 10 MEPs, expressed as a percentage of M_max, from one participant at each of the three cadences assessed (30, 60 and 90 rpm). In this example, MEPs were expressed as a percentage of M_max and were 27, 47.9 and 66.9% at 30, 60 and 90 rpm, respectively. As a group, MEP amplitudes were 36.1, 50.9 and 62.5% M_max at 30, 60 and 90 rpm, respectively (main effect for cadence: P < 0.001; Fig. 3A). Pairwise comparisons revealed that MEPs elicited at each cadence were significantly different from the other (P < 0.05 for all comparisons), with MEPs increasing as cadence increased.

Fig. 2.

Average motor evoked potential (MEP; A and C) and cervicomedullary evoked potential (CMEP; B and D) traces following 8 stimulations during arm cycling at 30 rpm (solid gray line), 60 rpm (dotted black line), and 90 rpm (solid black line) at the 6 o'clock (A and B) and 12 o'clock (C and D) positions. Amplitudes are expressed as a percentage of maximal M-wave (M_max).

Fig. 3.

Group data (means ± SE, n = 10) at the 6 o'clock position during the flexion phase for MEP amplitude (A), background EMG (bEMG) of the biceps brachii prior to transcranial magnetic stimulation (TMS; B), and bEMG of the triceps brachii prior to TMS (C), as well as group data (means ± SE, n = 7) at the 6 o'clock position for CMEP amplitude (D), bEMG of the biceps brachii prior to TMES (E), and bEMG of the triceps brachii prior to TMES (F). MEP and CMEP amplitudes are expressed relative to the M_max taken during cycling at the same cadence, and bEMG is expressed relative to the maximum EMG found during the 90 rpm trial. *Significant difference (P < 0.05) between cadences. #Differences that approached significance.

BEMG FOR MEPS.

A significant main effect for cadence was observed for the bEMG of the biceps brachii (main effect for cadence: P < 0.001; Fig. 3B) and triceps brachii (main effect for cadence: P = 0.035; Fig. 3C). For the biceps brachii, pairwise comparisons revealed that bEMG was significantly different at each cadence (P < 0.05 for all comparisons), whereas the bEMG was different only between 30 and 90 rpm (P < 0.05; Fig. 3C) for the triceps brachii.

Spinal excitability during elbow flexion.

CMEP AMPLITUDE.

Figure 2B shows the average of 10 CMEPs, expressed as a percentage of M_max, from one participant at each of the three cadences assessed (30, 60 and 90 rpm). In this example, CMEPs were expressed as a percentage of M_max and were 9, 12.5 and 19.2% at 30, 60 and 90 rpm, respectively. As a group, CMEP amplitudes were 13.8, 17.4 and 24.3% M_max at 30, 60 and 90 rpm, respectively (main effect for cadence: P < 0.01; Fig. 3D). Pairwise comparisons revealed CMEPs elicited at 90 rpm were significantly larger than at 30 (P = 0.002) and 60 (P = 0.013) rpm with a trend toward a significant difference between 30 and 60 rpm (P = 0.059).

BEMG FOR CMEPS.

A significant main effect for cadence was observed for the bEMG of the biceps brachii (P = 0.001; Fig. 3E) with pairwise comparisons indicating that 90 rpm produced significantly higher bEMG activity levels than both 30 (P = 0.002) and 60 (P = 0.038) rpm. The difference between 30 and 60 rpm was not significantly different, although there was a trend (P = 0.09). Although there was a pattern of increased activity as a function of cadence in the triceps brachii activity, the differences were not statistically different (main effect for cadence: P = 0.136; Fig. 3F).

BEMG OF BICEPS BRACHII AND TRICEPS BRACHII BETWEEN STIMULATION TYPES AS FUNCTION OF CYCLING CADENCE DURING ELBOW FLEXION.

There was no significant main effect for “stimulation type” (P = 0.93), nor was there an interaction of “stimulation type” and “cadence” (P = 0.72) for the bEMG of the biceps brachii. The bEMG of the triceps brachii activity was significantly higher during TMES compared with TMS (main effect for stimulation type: P = 0.018) with no interaction effect between “stimulation type and cadence” (P = 0.73). Thus general comparisons between changes in MEP and CMEP amplitudes are warranted.

Cadence-dependent Changes in CSE during the Extension Phase of Arm Cycling

CSE during elbow extension.

MEP AMPLITUDE.

Figure 2C shows the average of 10 MEPs, expressed as a percentage of M_max, from one participant at each of the three cadences assessed (30, 60 and 90 rpm). In this example, MEPs expressed as a percentage of M_max were 6.4, 6.1 and 10.2% at 30, 60 and 90 rpm, respectively. As a group, MEP amplitudes at the 12 o'clock position were 4.3, 4.7 and 9.1% M_max at 30, 60 and 90 rpm, respectively (main effect for cadence: P = 0.027; Fig. 4A). Pairwise comparisons revealed MEPs at 90 rpm were significantly greater than those recorded at 30 and 60 rpm (P = 0.036 and 0.016, respectively). There was no difference in MEP amplitudes between 30 and 60 rpm (P = 0.514).

Fig. 4.

Group data (means ± SE, n = 10) at the 12 o'clock position during the extension phase for MEP amplitude (A), bEMG of the biceps brachii prior to TMS (B), and bEMG of the triceps brachii prior to TMS (C), as well as group data (means ± SE, n = 7) at the 12 o'clock position for CMEP amplitude (D), bEMG of the biceps brachii prior to TMES (E), and bEMG of the triceps brachii prior to TMES (F). MEP and CMEP amplitudes are expressed relative to the M_max taken during cycling at the same cadence, and bEMG is expressed relative to the maximum EMG found during the 90 rpm trial. *Significant difference (P < 0.05) between cadences.

BEMG FOR MEPS.

There was no significant main effect for cadence observed for the bEMG of the biceps brachii or triceps brachii (main effects for cadence: P = 0.267 and 0.053, respectively; Figs. 4, B and C).

Spinal excitability during elbow extension.

CMEP AMPLITUDE.

Figure 2D shows the average of 10 CMEPs, expressed as a percentage of M_max, from one participant at each of the three cadences assessed (30, 60 and 90 rpm). In this example, CMEPs expressed as a percentage of M_max were 5.6, 2.8 and 0.9% at 30, 60 and 90 rpm, respectively. As a group, CMEP amplitudes were 2.5, 1.3 and 0.5% M_max at 30, 60 and 90 rpm, respectively (main effect for cadence: P = 0.01; Fig. 4D). CMEPs elicited at each cadence were significantly different from the other (P < 0.05 for all comparisons), with CMEPs decreasing as cadence increased.

BEMG FOR CMEPS.

There was no significant main effect for cadence observed for the bEMG of the biceps brachii (P = 0.293; Fig. 4E). A significant main effect for cadence was observed for the bEMG of the triceps brachii (P = 0.012; Fig. 4F), with pairwise comparisons indicating that triceps brachii activity was significantly lower at 60 rpm compared with both 30 and 90 rpm (P = 0.046 and 0.005, respectively). There was no difference in bEMG between 30 and 90 rpm (P = 0.514).

BEMG OF BICEPS BRACHII AND TRICEPS BRACHII BETWEEN STIMULATION TYPES AS FUNCTION OF CYCLING CADENCE DURING ELBOW EXTENSION.

There was no significant main effect for “stimulation type” nor was there an interaction of “stimulation type and cadence” for either the biceps brachii (P = 0.8 and P = 0.51) or triceps brachii (P = 0.27 and P = 0.26) for bEMG.

DISCUSSION

This report is the first to demonstrate the influence of arm cycling cadence on CSE. In the present study, CSE projecting to the biceps brachii increased with increasing cycling cadence during elbow flexion, as demonstrated via increased MEP amplitudes (Figs. 2A and 3A). Changes in spinal excitability, as indicated via CMEP amplitudes, followed the same pattern as the MEPs, suggesting that enhanced spinal excitability could partially account for the increase in MEP amplitudes (Figs. 2B and 3D). Changes in CSE during elbow extension, however, were different from those that occurred during elbow flexion. Although CSE was higher at 90 rpm than during 30 and 60 rpm, as shown via a larger MEP amplitude (Fig. 4A), changes in spinal excitability as indicated via CMEP amplitudes did not follow the same pattern (Fig. 4D). CMEP amplitudes successively decreased in amplitude as cadence increased. Thus, as cadence increases, supraspinal excitability appears to be enhanced throughout arm cycling, whereas changes in spinal excitability are phase dependent.

Cadence-dependent Changes in CSE during Elbow Flexion

The increase in MEP amplitude during elbow flexion as cycling cadence increased likely represents an increase in voluntary effort, which is consistent with the pattern of increased bEMG of the biceps brachii (Fig. 3B). This increased supraspinal excitability with higher cadences is indirectly supported by the finding of Christensen et al. (2000), who examined whether changes in cerebral activation, as estimated via measurements of cerebral blood flow, occurred during leg cycling at various cadences and workloads. They demonstrated a significant positive correlation between motor cortex activation and the rate of leg cycling. This suggests that, as arm cycling cadence increased in the present study, so too did the level of activation of the motor cortex, possibly increasing the excitability of cortical neurons projecting to the spinal cord, leading to increased responsiveness to TMS.

Changes in MEP amplitude, however, can result from changes in neuronal excitability at supraspinal and/or spinal locations. Using TMS and TMES, enhanced excitability of the corticospinal pathway projecting to the biceps brachii has been well-documented during different intensities of isometric contractions of the elbow flexors (Martin et al. 2006; Pearcey et al. 2014), with changes in both supraspinal and spinal excitability being contributing factors. Given that the pattern of change in MEP and CMEP amplitudes was similar during elbow flexion as cadence increased (Fig. 3, A and D), the increase in MEP amplitude is partially generated via increased spinal excitability. Higher effort levels, consistent with increased levels in biceps brachii bEMG, would presumably be met by increases in the recruitment and/or firing rates of spinal motoneurons as the cadence of arm cycling increased. Whether the enhanced spinal excitability was due to changes in the intrinsic electrical properties of spinal motoneurons and/or changes in the balance of excitatory and inhibitory input to the motoneuron pool is unclear. Currently, there is no detailed information as to changes in motoneuron electrical properties during rhythmic and alternating motor outputs in humans, although we recently demonstrated that spinal excitability was higher during arm cycling than an intensity-matched tonic contraction (Forman et al. 2014). Desmedt and Godaux (1977), however, used motor unit recordings to demonstrate that increasing the velocity of a voluntary isometric muscle contraction lowered the threshold for motoneuron activation, presenting the possibility that a lowering of motoneuron threshold as cycling cadence increased enhanced motoneuron excitability, thus increasing the gain of the motoneuron pool.

Work from the adult decerebrate cat during fictive locomotion and scratch (CPG-mediated motor outputs), however, suggests that changes in motoneuron properties may not depend on the intensity of the motor output (Krawitz et al. 2001; Power et al. 2010). For example, hyperpolarization of the voltage threshold for action potential initiation and reduced afterhyperpolarization amplitude were evident prior to and throughout the motor output. It is also worth noting that: 1) changes in motoneuron properties are similar between fictive locomotion and scratch, even though scratch is typically a faster motor output (Krawitz et al. 2001; Power et al. 2010); and 2) the degree of Vth hyperpolarization was not correlated with the amplitude of the locomotor or scratch drive potentials (Krawitz et al. 2001; Power et al. 2010), which provide a general index of the drive to the motoneuron pool (Jordan 1983). In fact, during fictive locomotion in the adult decerebrate rat, we recently demonstrated that the modulation of motoneuron excitability occurs across the motor pool and is not restricted to motoneurons engaged in locomotion (MacDonell et al. 2015). Finally, computer modeling of spinal motoneuron properties during fictive locomotion suggests that the rate of membrane depolarization is unlikely to be the main determinant of voltage threshold modulation (Dai 2000). Thus, although it is possible that an increase in the size of the TMES-evoked CMEPs during elbow flexion in the present study as cadence increased was due to changes in spinal motoneuron properties, we consider this to be a remote possibility. Rather, we suggest that changes in synaptic input to the spinal motoneuron pool can account for the observed changes in CMEP amplitudes (see below).

Cadence-dependent Changes in CSE during Elbow Extension

Increased CSE was demonstrated during elbow extension as cadence increased while spinal excitability decreased (Fig. 4, A and D). This suggests that the increase in MEP amplitude was due to enhanced excitability at the supraspinal level. Interestingly, this occurred without a concomitant increase in biceps brachii bEMG (Fig. 4B). Enhanced excitability of the cortex projecting to the biceps brachii motor pool during extension as the cadence increases may be due to cortical “spread” from both homologous and heterologous cortical connections. Bilateral arm cycling was used in the present study, which means that, when the biceps brachii of the dominant arm (i.e., arm from which recordings were made) was engaged in elbow extension, the contralateral limb was producing elbow flexion and thus likely had an increase in biceps brachii activity. It is possible that the contralateral motor cortex projecting to the biceps brachii of the dominant arm during elbow extension received excitatory input from the ipsilateral motor cortex projecting to the nondominant arm that was simultaneously involved in elbow flexion. Interhemispheric connections between homologous motor cortical areas have been previously suggested in studies examining central fatigue (Takahashi et al. 2011).

Potential Mechanisms Underlying Phase- and Cadence-dependent Changes in Spinal Excitability

It is likely that differences in synaptic input to the motoneuron pool, such as changes in reciprocal inhibition and/or presynaptic inhibition, play a significant role in the modulation of spinal excitability during arm cycling, as opposed to changes in the intrinsic properties of spinal motoneurons, as previously discussed. During fictive scratch in the turtle preparation, for example, excitation and inhibition to the motoneuron occur simultaneously with one being larger than the other, a process referred to as balanced networks (Berg et al. 2007). Once in a balanced state, changes in excitation and inhibition to the motoneuron pool have been suggested to operate in a push-pull fashion (Johnson et al. 2012). Applying the push-pull scheme to the present study would indicate that the motoneuron pool would be depolarized (increased gain) during elbow flexion by coupling an increase in excitatory drive to a decrease in inhibitory input (i.e., disinhibition) with hyperpolarization (decreased gain) during elbow extension resulting from increased inhibitory input and decreased excitation (i.e., disfacilitation) (Johnson et al. 2012). A likely candidate known to heavily regulate motoneuron output during rhythmic and alternating motor outputs is reciprocal inhibition. Reciprocal inhibition from the tibialis anterior to the soleus is reduced during the stance phase of locomotion when the soleus is active and increased during the swing phase when the tibialis anterior is active and the soleus is inactive (Petersen et al. 1999). Subsequent work demonstrated that reciprocal inhibition from the tibialis anterior to the soleus tended to decrease as cycling rate increased during soleus activation (Pyndt et al. 2003). Thus it is possible that disynaptic Ia reciprocal inhibition of the biceps brachii from the triceps brachii (Katz et al. 1991) would be reduced during elbow flexion, resulting in disinhibition of the motoneuron pool, and then be increased during elbow extension, thus decreasing motoneuron excitability in a push-pull fashion during arm cycling. This effect could be amplified by an increase in cadence, similar to the findings of Pyndt and colleagues (2003).

Modulation of afferent feedback is also a putative mechanism to explain the observed changes in spinal excitability as cycling cadence increased. For example, suppression of soleus H-reflex gain has been reported when running compared with walking (Capaday and Stein 1987; Ferris et al. 2001; Simonsen and Dyhre-Poulsen 1999), suggesting an increase in presynaptic inhibition. Similarly, the amplitudes of somatosensory evoked potentials and H-reflexes are suppressed as leg cycling cadence increases (Staines et al. 1997), an effect that is phase dependent (i.e., suppression when muscle is inactive and facilitation when muscle is active) (Larsen and Voigt 2004; Sakamoto et al. 2004). Changes in supraspinal processing and/or spinal reflex gain may be functionally important by preventing motoneuron saturation, thus allowing the spinal motoneuron pool to effectively respond to altered central or afferent input (Capaday and Stein 1987). It is thus possible that, as the cycling cadence increased in the present study, there were changes in the afferent feedback to the spinal motoneuron pool during both elbow flexion and extension (reduced excitation in combination with increased reciprocal inhibition).

Methodological Considerations

There are several factors to consider in the interpretation of the present results.

First, we do not have a measure of cadence variability; however, participants were able to maintain the selected cadences quite well. Second, as demonstrated in previous work (Marsh and Martin 1995), increased pedaling rates cause a shift to earlier activation and peak EMG (see biceps brachii in Fig. 1). Thus, in the present study, TMS and TMES were elicited at slightly different portions of the EMG burst during the different cadences (see Fig. 1B). We chose to elicit MEPs and CMEPs at a given position (6 and 12 o'clock) to ensure that the joint angle was the same, given the strong influence of joint angle on CSE, especially with respect to reciprocal inhibition at the spinal level (Hyngstrom et al. 2007). We are currently unsure if this may have affected the recorded evoked potentials. Third, the “intensity” of cycling can be modified by changing the cadence and/or the mechanical load. In the present study, because power output was kept constant as cadence increased, the torque required to rotate the arm crank was reduced as the cadence increased from 30 to 90 rpm. Even with a reduced torque, however, biceps brachii EMG activity increased during elbow flexion and did not change during elbow extension. It would perhaps be expected that a reduced torque would be associated with less biceps brachii EMG during the elbow flexion phase of arm cycling, which was not the case (Fig. 3B). The relationship between EMG and muscle output during cyclical movements, however, is not linear, as the movement intensity is altered. For example, increases in EMG can occur in the absence of changes in power output during leg cycling (Wakeling et al. 2010). Additional work has demonstrated that the neural control of force output and speed during locomotor activities may be differentially controlled in both humans and animals (Christensen et al. 2000; Hundza and Zehr 2009; Larsen et al. 2006; Shik et al. 1966). This possibility remains to be examined as it relates to CSE during arm cycling. Finally, it is presently unclear whether the increase in CSE demonstrated during the elbow flexion phase of arm cycling depends on the cadence, or whether it is a general response of the motor system as motor output intensity increases. Our laboratory's previous work, however, demonstrated that supraspinal excitability was higher during the elbow flexion phase of arm cycling compared with an intensity-matched tonic contraction (Forman et al. 2014). Whether this difference in supraspinal excitability between these two motor outputs is maintained or altered (increased or decreased) as motor output intensity increases (cycling cadence vs. tonic contraction intensity) remains to be examined.

Conclusion

We demonstrated increases in CSE of the biceps brachii during the elbow flexion and extension phases of arm cycling as cadence increased. While increased spinal excitability could partially account for the overall enhancement of CSE during elbow flexion, this was not the case during elbow extension. In fact, during the elbow extension phase of arm cycling, spinal excitability decreased as cycling cadence increased. The phase- and cadence-dependent modulation of spinal excitability is likely mediated, at least in part, by changes in synaptic input (e.g., reciprocal inhibition and/or afferent input) to the spinal motoneuron pool. This possibility remains to be examined.

DISCLOSURES

No conflicts of interest, financial or otherwise, are declared by the author(s).

AUTHOR CONTRIBUTIONS

Author contributions: D.A.F., D.C.B., and K.E.P. conception and design of research; D.A.F. and D.T.G.P. performed experiments; D.A.F. and D.T.G.P. analyzed data; D.A.F., D.T.G.P., D.C.B., and K.E.P. interpreted results of experiments; D.A.F., D.C.B., and K.E.P. prepared figures; D.A.F., D.C.B., and K.E.P. drafted manuscript; D.A.F., D.C.B., and K.E.P. edited and revised manuscript; D.A.F., D.T.G.P., D.C.B., and K.E.P. approved final version of manuscript.