Journal of Electromyography and Kinesiology 25 (2015) 355–362

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Superficial shoulder muscle co-activations during lifting tasks: Influence of lifting height, weight and phase Y. Blache a,⇑, F. Dal Maso a, L. Desmoulins a, A. Plamondon b, M. Begon a,c a

Laboratoire de Simulation et Modélisation du Mouvement, Département de Kinésiologie, Université de Montréal, Québec, Canada Institut de recherche Robert-Sauvé en santé et en sécurité du travail (IRSST), Montréal, Québec, Canada c Centre de Recherche du Centre Hospitalier Universitaire Sainte-Justine, Montreal, Québec, Canada b

a r t i c l e

i n f o

Article history: Received 4 August 2014 Received in revised form 22 October 2014 Accepted 12 November 2014

Keywords: Muscle focus Electromyography Co-contraction Glenohumeral joint Moment arm

a b s t r a c t This study aimed to assess the level of co-activation of the superficial shoulder muscles during lifting movement. Boxes containing three different loads (6, 12, and 18 kg) were lifted by fourteen subjects from the waist to shoulder or eye level. The 3D kinematics and electromyograms of the three deltoids, latissimus dorsi and pectoralis major were recorded. A musculoskeletal model was used to determine direction of the moment arm of these muscles. Finally an index of muscle co-activation named the muscle focus was used to evaluate the effects of lifting height, weight lifted and phase (pulling, lifting and dropping phases) on superficial shoulder muscle coactivation. The muscle focus was lower (more co-contraction) during the dropping phase compared to the two other phases (13%, p < 0.001). This was explained by greater muscle activations and by a change in the direction of the muscle moment arm as a function of glenohumeral joint position. Consequently, the function of the shoulder superficial muscles varied with respect to the glenohumeral joint position. To increase the superficial muscle coactivation during the dropping phase may be a solution to increase glenohumeral joint stiffness. Ó 2014 Elsevier Ltd. All rights reserved.

1. Introduction The glenohumeral joint has the greatest articular mobility in the human body (Halder et al., 2001) with its six degrees of freedom (three rotations and three translations). However this mobility is at the expense of its stability (Veeger and van der Helm, 2007). Since the shapes of the humeral head and the glenoid fossa do not ensure a complete congruence, joint stability is partially ensured by the capacity of the muscles to increase joint stiffness (Granata and Gottipati, 2008). Glenohumeral joint stiffness is mainly provided by the rotator cuff muscles (i.e. supraspinatus, infraspinatus, subscapularis, and teres minor) (Lee et al., 2000) to limit humerus head translations (Escamilla et al., 2009; Sharkey and Marder, 1995). By contrast the main function of the superficial muscles inserted on the humerus shaft, such as the deltoids, pectoralis major, and latissimus dorsi, is to produce force to move the upper limb. However, some studies have pointed out that superficial muscles may also contribute to glenohumeral joint stiffness if they are activated with antagonistic efforts (Kido et al., 2003; Veeger and van der Helm, 2007). Consequently, in movements involving glenohumeral rotations, shoulder muscle coordination ⇑ Corresponding author. Tel.: +1 514 343 6111x44017. E-mail address: [email protected] (Y. Blache). http://dx.doi.org/10.1016/j.jelekin.2014.11.004 1050-6411/Ó 2014 Elsevier Ltd. All rights reserved.

should produce a trade-off between force production to generate joint torque and maintaining glenohumeral joint stiffness (Veeger and van der Helm, 2007). Joint stiffness is increased by the co-activation (i.e. the simultaneous activation of agonist and antagonist muscles) of the muscles crossing the joint (Basmajian and DeLuca, 1985; Hogan, 1980; Morgan et al., 1978; Stokes and Gardner-Morse, 2003). The co-contraction index, which is based on agonist/antagonist joint moment and electromyography, is usually calculated to reflect joint stiffness during multi-joint dynamic exercises (Kellis et al., 2003). This index has been mainly applied to knee (Patsika et al., 2014; Rao et al., 2009) and elbow joints (Song et al., 2013). In both of these joints, the definition of agonist and antagonist muscles is obvious since the moments produced by muscle pairs are in an opposite direction. Consequently, the equation of Falconer and Winter (1985) that quantifies the co-contraction index is relevant to reflect joint stiffness. However, for ball and socket joints, the moment arms of all the muscles that cross the joint are not strictly opposed. It therefore becomes more complicated to define agonist/ antagonist pairs of muscles at the glenohumeral joint especially as the orientation of the muscles changes during arm rotation. Another index, referred to the muscle focus (MF), may be more appropriate to assess the co-contraction of the muscles surrounding the glenohumeral joint. MF was developed to assess muscle

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selectivity, i.e. the capacity to activate only muscles that contribute to movement (Yao et al., 2004). It is based on the electromyography (EMG) recordings as well as the direction of the muscle moment arms that derive from a musculoskeletal model. MF value ranges between 0 and 1; the lower the MF, the greater the muscle co-contractions, meaning that the activated muscles act in opposite directions. Thus, MF enables to evaluate the resultant of the forces produced by a given set of muscles inserted on the same bone and acting around a common joint. Among daily life activities, lifting tasks mostly involve the glenohumeral joint where kinematics and shoulder muscle activation vary with height and load (Anton et al., 2005; Yoon et al., 2012). A high and heavy lifting leads to a greater shoulder muscle activation and may also involve glenohumeral instability which are factors that reinforce shoulder injury risks. Therefore, we may wonder whether shoulder muscle co-contractions are influenced by lifting height, weight lifted, and task phase (pulling vs. lifting vs. dropping phases) to handle change in glenohumeral joint stiffness. Besides, although it is well known that rotator cuff muscles contribute to increasing glenohumeral joint stiffness (Escamilla et al., 2009), some questions remain concerning the contribution of superficial muscles to glenohumeral joint stiffness. To the best of our knowledge, no study has evaluated co-contractions of the superficial shoulder muscles with insertion on the humerus that drive glenohumeral kinematics during lifting tasks. Hence, this study aimed to determine the level of co-activation of superficial shoulder muscles during lifting movements using MF; more specifically the effect of lifting height, weight lifted and movement phase on muscle co-activation was assessed. According to the MF definition, the studied muscles have to respect three conditions (i) be involved in glenohumeral movement (ii) be inserted on the humerus bone, and (iii) their activation must be able to be measured by EMG. Consequently, the five superficial muscles taken into consideration for this study included the anterior deltoid, middle deltoid, posterior deltoid, pectoralis major, and the superior head of the latissimus dorsi. The hypothesis was that increased lifting height and/or weight, and the last phase of the movement (dropping phase) must lead to a lower MF because of (i) the greater muscle activation and (ii) higher antagonistic action of the shoulder muscles. 2. Methods 2.1. Participants Fourteen healthy male subjects volunteered in this study (mean ± SD: age, 26.1 ± 1.32 years; height, 1.80 ± 0.04 m; mass, 75.2 ± 8.82 kg). They provided a written informed consent. The protocol was approved by the University Ethics Committee (N°11-068-CERSS-D). None of the participants presented current or previous shoulder, elbow, or wrist injury. 2.2. Instrumentation and data collection Only the right side of each participant was analyzed, assuming that the right and left sides of the upper body behaved symmetrically (Nielsen et al., 1998). In accordance to a previous kinematic shoulder model (Jackson et al., 2012) (Fig. 1), 25 reflective markers were placed on the skin of the thorax (xiphoid process, 3 markers on the manubrium, 1st and 10th thoracic vertebrae), on the right side of the clavicle (sterno-clavicular joint, acromio-clavicular joint), scapula (acromion tip, acromial angle, inferior angle, trigonum spinae, superior angle), humerus (lateral and medial epicondyles), forearm (ulnar and radial styloid process), hand (proximal part of the 2nd and 3rd metacarpus, distal part of the 2nd and

Fig. 1. Placement of the reflective markers in line with the model of Jackson et al. (2012).

5th metacarpus) and the box (four superior angles). Each trial was recorded using an 18-camera Vicon™ motion analysis system at 200 Hz (Oxford Metrics Ltd., Oxford, UK). The electromyograms measurements (EMG) of the anterior deltoid, middle deltoid, posterior deltoid, pectoralis major, and latissimus dorsi superior head muscles were taken with pairs of wireless surface electrodes at 1000 Hz (Trigno, Delsys Inc., Boston, MA). After shaving and cleaning the skin with alcohol, electrodes were positioned on the belly of the muscles according to the SENIAM (Surface ElectroMyoGraphy for the Non-Invasive Assessment of Muscles) recommendations for electrode locations (Hermens et al., 2000). Kinematic and EMG signals were synchronized using the Nexus 1.8.2 software (Vicon, Oxford, UK). The experimental tasks consisted in lifting a box between shelves. The size of the box was 0.08  0.395  0.345 m in height, width, and length respectively. To facilitate and standardize the grip of the box and ensure symmetrical movement, two handgrips were positioned on the right and left sides of the box.

2.3. Experimental procedures Prior to the tests, isometric maximal voluntary contractions for five muscles (anterior deltoid, middle deltoid, posterior deltoid, pectoralis major, and latissimus dorsi superior head) were randomly performed by the participant according to Ekstrom et al. (2005) and Boettcher et al. (2008) instructions (Table 1) to elicit maximum muscle activation. The participants had to exert maximum force against an experimenter during five seconds. Verbal encouragement was provided throughout the duration of the maximum effort. Two trials per muscle were performed. The rest interval was 30 s between repetitions and 60 s between trials for different muscles. The same experimenter performed all the testing sessions to reduce inter-subject variability in segment position and resistance. To familiarize themselves with the procedure, participants performed a total of 24 sagittal lifting movements at different heights and with the three box masses. The experimental test consisted in lifting a box positioned on a shelf from hip level to a shelf located

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Table 1 Description of the isometric maximal voluntary contractions tests for anterior deltoid, middle deltoid, posterior deltoid, latissimus dorsi and pectoralis major muscles. Muscle

Position

Instructions

Anterior deltoid Middle deltoid Posterior deltoid

Seated, shoulder flexion 90° Seated, shoulder abduction 90° Prone, shoulder extension 90°

Latissimus dorsi

Seated, shoulder adduction 90°

Pectoralis major

Seated, shoulder adduction 30°

Shoulder flexion 90° with elbow extended. Resistance applied at the elbow Shoulder abduction 90°, hand in a prone position. Resistance applied the elbow Shoulder horizontal abduction 90°, elbow flexed 90° with thumb pointing at ceiling. Resistance applied at the elbow Arm adduction at 90° abduction with elbow flexed 90°. Resistance applied at the elbow Arm adduction at 30° abduction with elbow flexed 90°. Resistance applied at the elbow

either at shoulder level (H1) or eye level (H2). Thus, the vertical distances between the initial shelf and H1 or between the initial shelf and H2 were on average equal to 51.5 ± 4.9 cm and 79.5 ± 4.8 cm respectively. For each height, three boxes of different masses were lifted: 6 kg, 12 kg, and 18 kg. Consequently six conditions were tested (2 heights  3 loads). The six conditions were performed in random order with alternating 30-s rest periods. In each condition, three repetitions were collected with alternating 3-min rest periods. No instruction concerning the velocity of the movement and the handgrip technique was given to the participant. The participants chose their preferred horizontal distance from the shelves and were instructed to minimize feet displacements. One back step was allowed. 2.4. Data processing Based on the frequency and residual analysis of the raw signals (Winter, 2009), kinematic raw data were filtered using a zero-lag 4th-order Butterworth low-pass filter at 8 Hz. EMG raw signals were band-pass filtered using a zero-lag 4th-order Butterworth band-pass filter between 15 and 500 Hz (De Luca et al., 2010) before calculating Root Mean Squared (RMS) EMG using a 250ms sliding window (Merletti, 1999). Then, RMS EMG of each muscle obtained for each lifting condition was normalized with respect to the maximum RMS EMG value obtained during the two trials of isometric maximal voluntary contractions. For each trial, the movement was decomposed into three phases, namely the pulling phase (P1; 0–20% of the total movement duration), lifting phase (P2; 21–60%), and dropping phase (P3; 61–100%).

actuated by 57 muscle–tendon units including the 5 muscles of interest (Fig. 2). The generic model was scaled to match subjects’ anthropometry using the experimental positions of the skin markers. An inverse kinematic procedure was used to obtain joint angles by minimizing the Euclidean distance between the positions of the experimental markers and the positions of the virtual markers of the model at each sample time. 2.6. Muscle co-activation variables The direction of the force vector exerted by the anterior deltoid, middle deltoid, posterior deltoid, pectoralis major, and latissimus dorsi superior head on the humerus were calculated using an OpenSim plugin (van Arkel et al., 2013). Briefly, the muscles attached to the humerus were identified; the path of each muscle was retrieved including wrapping objects and via points. The coordinates of the line of actions were then expressed in the glenohumeral joint coordinate system (Fig. 3). Finally, the unit vectors of the moment arm of each muscle with respect to the glenohumeral joint center were calculated at each sample time. Since pectoralis major and latissimus dorsi were modelled by three lines of action in the musculoskeletal model, the line of action that best corresponded to the position of the EMG electrodes was used to compute MF according to Yao et al. (2004):

,   X X 5 5   ! MFðtÞ ¼  dm ðtÞ  EMGnm ðtÞ EMGnm ðtÞ;  m¼5 m¼5

ð1Þ

where at each sample time (t), EMGn was the normalized EMG of the five muscles (m: anterior deltoid, middle deltoid, posterior

2.5. Musculoskeletal model An OpenSim generic model of the shoulder, based on the complementarity of three existing models (Holzbaur et al., 2005; Nikooyan et al., 2011; Steele et al., 2013), was implemented using the Graphical User Interface of OpenSim (Delp et al., 2007). The model was composed of 9 segments (ground, thorax + spine, right clavicle, right scapula, right humerus, right ulna, right radius, right hand, and box) and 22 degrees of freedom. The joint between the ground and the thorax included six degrees of freedom (three translations and three rotations). Sternoclavicular, acromioclavicular, and glenohumeral joint movements were defined as balland-socket joints (three rotational degrees of freedom). The axial rotation of the clavicle was limited to 20° (Nikooyan et al., 2011). Elbow and wrist movements were modelled using two degrees of freedom: flexion/extension and prono-supination for the elbow; flexion/extension and deviation for the wrist. It was assumed that the contact point between the handgrip and the hand remained fixed throughout the trial: it corresponded to the capitate bone for the hand and the middle of the right handgrip for the box. Thus, the joint between the hand and the box was defined by three rotational degrees of freedom. The musculoskeletal model was

Fig. 2. OpenSim musculoskeletal model of the upper limb with the 57 actuator muscle–tendon units (including the five muscles of interest) and the box.

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Fig. 3. Lateral view of the right humerus with the direction of the force vector exerted by the anterior deltoid, middle deltoid, posterior deltoid, pectoralis major, and latissimus dorsi superior head at the beginning of the movement. The direction of the force vectors take into consideration the wrapping objects and via point, explaining why some insertions are not attached to the humerus.

! deltoid, pectoralis major, and latissimus dorsi superior head); dm corresponded to the unit vector of the moment arm of each muscle. Values of MF range between 0 and 1. MF close to 1 indicates that activated muscles have similar functions and that consequently the level of co-contraction is low. In contrast, a MF close to zero indicates that activated muscles have opposite actions which corresponds to a high level of co-contraction. Finally, for each muscle pair (10 combinations for 5 muscles), the scalar product of their moment arm was calculated at each sample time. A scalar product close to 1 indicates that the pair of muscles acts synergistically, while a scalar product close to 1 means that the pair of muscles acts antagonistically. 2.7. Statistics After testing the normality of the data with the Shapiro–Wilk test, the effects of lifting height, weight lifted, and phase (independent variables) on mean MF, mean scalar product of the pairs of muscles moment arm, and mean normalized EMGs (dependant variables) were successively assessed using a three-way repeated measures ANOVA. As the number of dependant variables was equal to 16 (1 MF, 10 scalar products, and 5 EMGs), the level of significance of each ANOVA was adjusted by Bonferroni correction yielding a significant p < 0.003 (0.05/16 dependant variables). When a significant F-value was found, post hoc pairwise comparisons were performed (Tukey post hoc) with a significant p-value set at 0.05.

Fig. 4. Time histories of the muscle focus (mean ± SD with the line and area, n = 14), for the conditions 6 kg (top), 12 kg (middle), and 18 kg (below). H1 corresponds to the shoulder level and H2 to the eye level.

3. Results 3.1. Muscle focus Our hypothesis was partially validated; a significant effect of the phase on MF was observed while no effect related to height and weight lifted was noted.

The changes in muscle focus as a function of time for the six conditions are presented in Fig. 4. The results of the three-way

Y. Blache et al. / Journal of Electromyography and Kinesiology 25 (2015) 355–362

ANOVA indicated no interaction effects of the lifting height, weight lifted, and phase on MF (F4,221 = 1.11, p = 0.35). The main effects indicated that the phase had a significant effect on MF (p < 0.001), which was about 13% lower during the dropping phase when compared to the pulling and lifting phases (Fig. 5). No effect of the mass and height on MF was pointed out, even if MF tended to decrease with the increase in lifting height (0.55 ± 0.13 versus 0.53 ± 0.14, for H1 and H2 respectively, p = 0.06). As many comparisons were performed, only the comparisons that enabled to explain the variation of MF were presented in the following sections.

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3.2. Mean normalized EMG The results of the three-way ANOVA on mean normalized EMG indicated independent effects of the lifting height (F1,211 = 3.63, p = 0.06), weight lifted (F2,221 = 48.2, p < 0.001), and phase (F2,221 = 72.2, p < 0.001) (Table 2). For each muscle, a significant increase in the mean EMG of over +40% was observed with weight enhancement from 6 kg to 18 kg. The increase in height led to a greater mean EMG of the middle deltoid only (+26%), even if a tendency to an enhancement of the mean EMG of the anterior deltoid was observed (+8.3%, p = 0.06). For all the muscles but the pectoralis major, the mean EMG was greater during the dropping phase compared to the pulling phase (between +40% and +158%). For the middle deltoid, posterior deltoid and latissimus dorsi, the mean EMG was also increased between the lifting and dropping phases (between +17% and +83%). 3.3. Muscle line of action and scalar product A qualitative analysis showed that the direction of the force vector exerted by the anterior deltoid, middle deltoid, posterior deltoid, pectoralis major, and latissimus dorsi superior head on the humerus changed during the lifting movement whatever the condition (see supplementary file, muscle_force_vector_direction.avi). The results of the three-way ANOVA on scalar product of muscle pairs showed independent effects of the lifting height (F2,221 = 0.79, p = 0.70), weight lifted (F2,221 = 3.04, p = 0.05), and phase (F2,221 = 17.8, p < 0.001) (Table 3). Weight lifted influenced only the scalar product of the muscle pair of the middle deltoid/ posterior deltoids, which was lower during the 6 kg conditions compared to the 12 or 18 kg conditions. Concerning the comparison between H1 and H2, a higher height (H2) led to a smaller scalar product of the anterior deltoid/posterior deltoid, posterior deltoid/ latissimus, and posterior deltoid/pectoralis major muscle pairs (2.7%, 8.9%, and 10%, respectively). Inversely, the scalar product of the muscle pair anterior deltoid/latissimus was greater (+24%) in H2 compared to H1. When comparing Phase 1 to Phase 3, five scalar products of muscle pairs decreased, two increased, and three remained unchanged (Table 3). From Phase 2 to Phase 3, five scalar product muscle pairs decreased, four increased, and one remained unchanged (Table 3). 4. Discussion

Fig. 5. Effects of the weight lifted (top), phase (middle), and height (bottom) on muscle focus. Phases 1 to 3 corresponded to pulling, lifting, and dropping respectively. H1 denoted shoulder level and H2 eye level.

The purpose of this study was to determine the level of co-activation of five superficial muscles (deltoids, pectoralis major, and latissimus dorsi) during lifting tasks to estimate their contribution to glenohumeral stability. To that aim, MF, an index of muscle co-activation, was calculated as a function of two lifting heights (shoulder versus eye level), three loads (6, 12, and 18 kg), and three phases of movement (pulling, lifting, dropping phases). Our hypothesis was partially supported. We observed that muscle co-activation was lower during the dropping phase compared to the pulling and lifting phases and that muscle co-activation tended to be lower when the box was lifted to eye level compared to shoulder level. In contrast, the weight lifted had no effect on superficial shoulder muscle co-activation. Shoulder muscle co-activation was influenced by the lifting phase. Indeed, the dropping phase led to a smaller MF compared to the lifting and pulling phase. This can be explained by an increase, during the dropping phase, in the mean EMG amplitude of the middle, posterior deltoids, and the latissimus dorsi (compared to pulling and lifting phases), as well as an increase in the anterior deltoid EMG (compared to the pulling phase). In addition, more than the half of the scalar products of the muscle pairs were

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Table 2 Mean ± standard deviation of normalized mean EMG of the anterior, middle and posterior deltoid, pectoralis major and latissimus dorsi muscles as a function of lifting weight, height (H1: shoulder level, H2: eye level) and phase (P1: pulling phase, P2: lifting phase, P3: dropping phase). Concerning the post hoc results, for each EMG muscle and each independent variable (weight, height and phase), means (in vertical reading) that do not share the same subscript (a, b or c) differ at p < 0.05. EMG ant. deltoid

EMG mid. deltoid

EMG post. deltoid

EMG pect. major

EMG latissimus

Weight 6 kg 12 kg 18 kg

0.18 ± 0.10a 0.28 ± 0.14b 0.31 ± 0.13b

0.17 ± 0.11a 0.24 ± 0.14b 0.24 ± 0.14c

0.05 ± 0.04a 0.09 ± 0.07b 0.08 ± 0.07b

0.07 ± 0.05a 0.14 ± 0.11b 0.19 ± 0.12c

0.05 ± 0.04a 0.07 ± 0.05b 0.07 ± 0.05b

Height H1 H2

0.24 ± 0.13a 0.26 ± 0.15a

0.19 ± 0.13a 0.24 ± 0.13b

0.07 ± 0.06a 0.08 ± 0.07a

0.13 ± 0.11a 0.11 ± 0.10a

0.06 ± 0.04a 0.06 ± 0.05a

Phase P1 P2 P3

0.15 ± 0.08a 0.30 ± 0.13b 0.30 ± 0.15b

0.12 ± 0.07a 0.22 ± 0.11b 0.31 ± 0.14c

0.05 ± 0.04a 0.06 ± 0.05b 0.11 ± 0.08c

0.11 ± 0.08a 0.16 ± 0.11b 0.11 ± 0.11a

0.05 ± 0.04a 0.06 ± 0.05b 0.07 ± 0.05c

Table 3 Mean ± standard deviation of scalar product of the 10 muscle pairs as a function of lifting weight, height (H1: shoulder level, H2: eye level) and phase (P1: pulling phase, P2: lifting phase, P3: dropping phase). Concerning the post hoc results, for each scalar product and each independent variable (weight, height and phase), means (in vertical reading) that do not share the same subscript (a, b or c) differ at p < 0.05. ant. deltoid/ mid. deltoid

post. deltoid/ latisimus

post. deltoid/ pect. major

Latissimus/ pect. major

0.08 ± 0.32a 0.06 ± 0.36a 0.02 ± 0.42a

0.76 ± 0.10a 0.73 ± 0.15a 0.76 ± 0.12a

0.52 ± 0.15a 0.55 ± 0.13a 0.51 ± 0.13a

0.12 ± 0.14a 0.10 ± 0.13a 0.11 ± 0.13a

0.86 ± 0.12a 0.84 ± 0.15a

0.05 ± 0.37a 0.06 ± 0.37a

0.78 ± 0.07a 0.71 ± 0.15b

0.50 ± 0.11a 0.55 ± 0.15b

0.10 ± 0.10a 0.12 ± 0.16a

0.84 ± 0.14a 0.81 ± 0.17a 0.90 ± 0.09b

0.09 ± 0.39a 0.03 ± 0.41b 0.10 ± 0.26a

0.77 ± 0.06a 0.80 ± 0.09a 0.66 ± 0.15b

0.53 ± 0.07a 0.47 ± 0.12b 0.58 ± 0.18c

0.09 ± 0.08a 0.09 ± 0.11a 0.15 ± 0.18b

ant. deltoid/ post. deltoid

ant. deltoid/ latissimus

ant. deltoid/ pect. major

mid. deltoid/ post. deltoid

mid. deltoid/ latissimus

Weight 6 kg 0.34 ± 0.32a 12 kg 0.29 ± 0.38a 18 kg 0.25 ± 0.48a

0.73 ± 0.08a 0.74 ± 0.09a 0.72 ± 0.09a

0.19 ± 0.10a 0.18 ± 0.13a 0.20 ± 0.10a

0.93 ± 0.05a 0.94 ± 0.03a 0.93 ± 0.03a

0.82 ± 0.18a 0.76 ± 0.25b 0.73 ± 0.31b

0.88 ± 0.11a 0.86 ± 0.13a 0.82 ± 0.18a

Height H1 0.29 ± 0.38a H2 0.29 ± 0.39a

0.72 ± 0.06a 0.74 ± 0.10b

0.21 ± 0.07a 0.16 ± 0.13b

0.92 ± 0.03a 0.93 ± 0.03a

0.79 ± 0.23a 0.74 ± 0.27a

Phase P1 P2 P3

0.70 ± 0.04a 0.68 ± 0.05a 0.80 ± 0.10b

0.17 ± 0.06a 0.22 ± 0.06b 0.16 ± 0.17a

0.94 ± 0.02a 0.92 ± 0.03b 0.92 ± 0.04b

0.80 ± 0.30a 0.71 ± 0.26b 0.79 ± 0.17a

0.29 ± 0.43a 0.19 ± 0.43b 0.39 ± 0.24c

significantly lower during the dropping phase compared to the pulling or lifting phases. Particularly, the actions of the anterior and posterior deltoids were more antagonistic during the dropping phase compared to the two other phases. Then, even if the deltoids muscles are included in the same muscle group, their line of action may very well have different functions with respect to glenohumeral joint rotation (supplementary file, muscle_force_vector_ direction.avi). In addition, the posterior deltoid and the latissimus dorsi, which were agonist at the beginning of the movement, did not quite act in the same direction at the end of the lifting movement. Consequently, the numerator showed less increase during the dropping phase than the denominator of the MF equation, resulting in a lower MF. These results are in accordance with some studies that pointed out a greater activation of the superficial muscles in tasks with high versus low arm elevations (Anton et al., 2001; Grieve and Dickerson, 2008). Besides, the changes in scalar product of muscle pairs may reflect the modification of the deltoids, pectoralis major, and latissimus dorsi functions during arm elevation (Favre et al., 2009; Giphart et al., 2013; Massimini et al., 2012). Finally, during the dropping phase, the subjects have to control the box speed and position. Yet, some studies (Gribble et al., 2003; Laursen et al., 1998) pointed out that a greater precision of movement involved an increase in the shoulder muscle activity to control said movement. A tendency (p = 0.06) of height effect on MF was observed. A decrease of MF during H2 compared to H1, may confirm that the greater the humerus elevation the greater the superficial shoulder muscles co-activation. Even if height effect is not significant in our study, this result is similar to those obtained by Granata and

mid. deltoid/ pect. major

Orishimo (2001) whereby for a task consisting in maintaining a load at different height, it was observed that the co-activation of the antagonist trunk flexor muscles increased with the height of the maintained load. No effect on MF was observed in regards to the weight lifted. The glenohumeral joint kinematics is probably minimally affected by the weight of the box resulting in similar scalar products of the muscles pairs (except for one of them). Thus even if the normalized EMGs were greater with load increase, the numerator and the denominator of the MF equation increased in the same proportion since no changes in scalar product were observed. This result does not confirm our hypothesis concerning the effect of weight lifted. Nevertheless, same outcomes were observed concerning the knee joint. Indeed, Rao et al. (2009) showed that the co-contraction index around the knee joint was not influenced by the load during squat movements. We may assume that the greater co-activation (i.e. smaller MF) of the superficial shoulder muscles reinforces the stiffness of the glenohumeral joint (Hogan, 1980, Morgan et al., 1978). This was especially the case during the dropping phase. Greater co-activation may improve glenohumeral stability; nevertheless, we were not able to establish if the changes in co-activation result from a modified glenohumeral joint stability. Indeed, muscle co-activation is dependent on joint stability as well as joint torque (Granata and Orishimo, 2001). As the glenohumeral joint torque is not constant throughout the phases and the conditions (height and weight), further studies are needed to assess the effect of glenohumeral joint instability on MF and superficial shoulder muscle co-activations.

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There are several limitations to this study. The first limit related to the accuracy of the clavicle and scapula kinematics due to soft tissue artifact (Leardini et al., 2005; Lempereur et al., 2010). As the inverse kinematics is performed with global optimization, the errors of the clavicle and scapula kinematics have an effect on the humerus kinematics and consequently the path of the muscles is affected. Nevertheless, shoulder kinematics remain challenging (Jackson et al., 2012) and no skin marker model has yet been validated while electromagnetic sensors are known to produce error up to 30° in arm axial rotation (Hamming et al., 2012). The second limit is that the MF did not include certain muscles such as the coracobrachialis, biceps brachialis (long and short heads), and triceps brachialis long head. Although these muscles may have an effect on glenohumeral joint stiffness, the activity of the coracobrachialis is not measurable with surface electrodes; the biceps and triceps muscles are bi-articular muscles and not directly connected to the humerus, therefore these muscles cannot be included in the MF equations that consider the humerus as a free segment. Consequently, advanced algorithms are needed to evaluate the co-activation of pluri-articular muscles crossing the glenohumeral joint during lifting tasks. 5. Conclusions In conclusion, the muscle focus is an index that deals with the complexity of the shoulder muscle path to determine agonist/ antagonist muscle effect on the glenohumeral joint. We found that the dropping phase increased the co-activation of the deltoids, pectoralis major, and latissimus dorsi. Elevation of the humerus at increased height decreased the muscle focus index due to an increase in the muscle activities and change in the direction of muscle action on the glenohumeral joint. These outcomes highlight that the superficial shoulder muscles have not the same function depending on the phase of the lifting movement, and that agonist/antagonist muscle pairs vary as a function of the glenohumeral joint position. It was assumed that the increase in these muscle co-activations improves glenohumeral joint stiffness. Conflict of Interest The authors state that there are no conflicts of interest. Acknowledgement We would like to thank the Institut National du Sport du Québec for the loan of the electromyographic wireless system. Appendix A. Supplementary material Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.jelekin.2014.11. 004. References Anton D, Shibley LD, Fethke NB, Hess J, Cook TM, Rosecrance J. The effect of overhead drilling position on shoulder moment and electromyography. Ergonomics 2001;44(5):489–501. Anton D, Rosecrance JC, Gerr F, Merlino LA, Cook TM. Effect of concrete block weight and wall height on electromyographic activity and heart rate of masons. Ergonomics 2005;48(10):1314–30. Basmajian JV, DeLuca CJ. Muscles alive. 5th ed. Baltimore, MD: Williams and Wilkins; 1985. Boettcher CE, Ginn KA, Cathers I. Standard maximum isometric voluntary contraction tests for normalizing shoulder muscle EMG. J Orthop Res 2008;26(12):1591–7. De Luca CJ, Gilmore LD, Kuznetsov M, Roy SH. Filtering the surface EMG signal: movement artifact and baseline noise contamination. J Biomech 2010;43(8):1573–9.

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Yoann Blache obtained his Ph.D degree in Human Movement Sciences (Biomechanics) from the University of Lyon (France) in 2011. He is currently a post-doctoral researcher in the Simulation & Movement Modeling laboratory at the University of Montréal (QC, Canada). His research interests include musculoskeletal modeling and shoulder function in manual handling tasks.

Fabien Dal Maso obtained his Ph.D. degree in Research Program in Sport and Human Movement Sciences (Neurosciences and Biomechanics) from the University of Toulouse III (Toulouse, France) in 2012. He is currently a post-doctoral researcher in the Simulation and Movement Modelisation laboratory at the University of Montreal, Quebec, Canada. His research interests are primarily in the area of the cerebral correlates of the muscle contraction, shoulder kinematics, muscle force estimation using electromyography-based models, and processing of the electroencephalography and electromyography signals.

Landry Desmoulins obtained his MSc in Engineering and Ergonomics of Human Movement from the University of the Mediterranean (Marseille, France) in 2012. He is a PhD candidate in the department of Kinesiology at the University of Montreal (Biomechanics). Current research interests concern occupational biomechanics and ergonomics with a focus on workplace shoulder injuries prevention.

André Plamondon received a B.Sc. and a M.Sc. in Sciences in Physical Activity from the University Laval, Quebec, Canada, and a Ph.D. in Sciences in Physical Activity from the University of Montreal, Quebec, Canada. He is currently researcher at the Occupational Health and Safety Research Institute Robert-Sauvé (IRSST), Montreal, Quebec, Canada. His research interest is focused on occupational biomechanics and mainly on the quantification of back loads during manual material handling tasks.

Mickaël Begon obtained his Ph.D. degree in Biomechanics and Bio-Engineering in 2006. He is currently Associate Professor in Kinesiology at the University of Montréal and researcher at St-Justine pediatric Hospital, Montréal (Québec), Canada. Mickaël Begon is head of the Simulation and Movement Modeling Team. His main research interests concern musculoskeletal modeling, computer simulation of human movement, and shoulder biomechanics.