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Effect of hammer mass on upper extremity joint moments
Applied Ergonomics 60 (2017) 231e239
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Applied Ergonomics journal homepage: www.elsevier.com/locate/apergo
Effect of hammer mass on upper extremity joint moments Nilanthy Balendra a, Joseph E. Langenderfer b, * a b
Department of Physical Therapy, Central Michigan University, Mount Pleasant, MI 48859, USA School of Engineering and Technology, Central Michigan University, Mount Pleasant, MI 48859, USA
a r t i c l e i n f o
a b s t r a c t
Article history: Received 3 February 2016 Received in revised form 29 November 2016 Accepted 2 December 2016
This study used an OpenSim inverse-dynamics musculoskeletal model scaled to subject-specific anthropometrics to calculate three-dimensional intersegmental moments at the shoulder, elbow and wrist while 10 subjects used 1 and 2 lb hammers to drive nails. Motion data were collected via an optoelectronic system and the interaction of the hammer with nails was recorded with a force plate. The larger hammer caused substantial increases (50e150%) in moments, although increases differed by joint, anatomical component, and significance of the effect. Moment increases were greater in cocking and strike/follow-through phases as opposed to swinging and may indicate greater potential for injury. Compared to shoulder, absolute increases in peak moments were smaller for elbow and wrist, but there was a trend toward larger relative increases for distal joints. Shoulder rotation, elbow varus-valgus and pronation-supination, and wrist radial-ulnar deviation and rotation demonstrated large relative moment increases. Trial and phase durations were greater for the larger hammer. Changes in moments and timing indicate greater loads on musculoskeletal tissues for an extended period with the larger hammer. Additionally, greater variability in timing with the larger hammer, particularly for cocking phase, suggests differences in control of the motion. Increased relative moments for distal joints may be particularly important for understanding disorders of the elbow and wrist associated with hammer use. © 2016 Published by Elsevier Ltd.
Keywords: Inverse-dynamics Kinetics
1. Introduction Throughout human history hammering has been a common manual task afforded by holding an object of relatively large concentrated mass in, or, by use of a handle, at some extended distance from the hand(s) (Kahrs et al., 2014). As the mass of the object is increased, and/or the object is moved further from the hand, the mass moment of inertia (MMOI) increases and the potential for the human to exert impulsive forces on the external environment and thereby perform work is consequently improved. However, the potential cost of this increased MMOI and associated work may be increased loads on the human body in the form of joint loadings and stresses applied to musculoskeletal tissues. In contrast to medieval times when a blacksmith formed metal into useful objects via repetitive hammering, modern industrial processes and the people that perform them have benefited from electromechanical automation of work. However, hammering remains an important task in carpentry and concrete form
* Corresponding author. E-mail address: [email protected] (J.E. Langenderfer). http://dx.doi.org/10.1016/j.apergo.2016.12.001 0003-6870/© 2016 Published by Elsevier Ltd.
construction, where it is the single most commonly performed work at 17% of the workday (Spielholz et al., 1998). Additionally, workers rate arm strain to be either high or very high in 52% of hammering tasks (Mattila et al., 1993). While the nature of hammering in terms of intensity and body position tends to be variable rather than fixed (Spielholz et al., 1998), by definition, hammer usage is extremely repetitive (Bernstein, 1967) and frequently occurs in a compromised overall posture which tends to cause cumulative fatigue related injuries (Mattila et al., 1993; Spielholz et al., 1998). Considerable attention has been devoted to investigating hammering from several perspectives. Alterations in hammer design in terms of mass and handle shape or length have been shown to affect many outcomes, including: ability to perform and complete hammering tasks (Konz, 1986; Schoenmarklin and Marras, 1989a,b), perceived upper extremity exertion (Karkowski et al., 2002/2003), arm muscle fatigue (Schoenmarklin and Marras, 1989a,b), body discomfort (Khan and Shakeb, 2009; Schoenmarklin and Marras, 1989a,b), and changes in grip strength before and after hammering (Knowlton and Gilbert, 1983). Additionally, when the orientation of the work surface is horizontal, rather than vertical, the human is able to attain better
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performance of the hammering task (Karkowski et al., 2002/2003, Schoenmarklin and Marras, 1989a,b) with decreased muscle exertion and discomfort (Schoenmarklin and Marras, 1989a,b). Previous investigations have observed wrist (Schoenmarklin and Marras, 1989a,b) and upper extremity (Cote et al., 2008, Cote et al., 2005) kinematics during actual (Cote et al., 2008, Cote et al., 2005; Schoenmarklin and Marras, 1989a,b) or simulated (Leventhal et al., 2010) hammering using electromechanical (Schoenmarklin and Marras, 1989a,b), optoelectronic (Cote et al., 2008, Cote et al., 2005) or imaging techniques (Leventhal et al., 2010). In general, angled hammer handles have been shown to reduce ulnar deviation during swinging phase (Knowlton and Gilbert, 1983; Schoenmarklin and Marras, 1989a,b), but increase radial deviation at the initial hammering position (Schoenmarklin and Marras, 1989a,b). When compared with other tasks, nailing with a hammer has been perceived to be more difficult following fatigue (Hammarskjold and Harms-Ringdahl, 1992). Conversely, when repetitive hammering has been used as a fatiguing activity, it has been shown to cause reduced elbow velocity, acceleration, range of motion and grip strength (Cote et al., 2008, Cote et al., 2005) while leaving wrist and shoulder kinematics and upper extremity EMG unaffected. When compared to hammering, tennis is not as common an activity among the general population in terms of frequency and duration of performance, but is similar in that a tool is used as an extension of the upper extremity in order to accomplish a manual task. Tennis racket mass and MMOI play an important role in the ability of players to control the racket during swing, ball strike and follow through and thus impart energy to a ball in order to return it to an opponent. Relatively small changes (10e15%) in racket mass and MMOI have been shown to substantially affect joint loads with calculated differences of 30e40% (Creveaux et al., 2013a,b; Rogowski et al., 2014). Hammers used in common carpentry tasks (455e900 g) generally have greater mass (100e200%) but smaller MMOIs (from 25 to 50%, measured in this study) compared to tennis rackets (Creveaux et al., 2013a,b). If one accepts that hammering is a relatively common activity among professional laborers and the general population, the question of upper extremity joint loading resulting from hammer usage becomes relevant. In spite of the ubiquitous nature of hammer use among lay persons and professional manual laborers, and the substantial number of previous investigations of hammering, estimates of the loads applied to upper extremity joints resulting from hammer tool usage are notably absent from scientific literature. Effects of hammer mass and MMOI on timing, joint kinematics and moments have been examined previously with only very brief and undetailed results (Putnam and Jenkins, 1993). In ecological psychology, previous studies of object manipulation have explored the relationship between intended object tool usage and resultant grasping location (Fitzpatrick et al., 2012; Wagman and Carello, 2003). Subjects asked to grasp objects in a manner preparatory to hammering do so in a way that minimizes the symmetry of object inertia relative to magnitude of overall inertia. As such, subjects are described as behaving in a way that minimizes diversity of muscle forces required to control motion relative to the amount of force required to move the object in order to generate an efficient hammering motion. However, as these studies explored human behavior and not motor control or biomechanics, any hypothesis on muscle forces is unsupported as little is known about how actual hammering affects humans in terms of timing, kinematics, or joint kinetics. An investigation to determine how different hammers affect upper extremity kinetics is a necessary initial step towards understanding the underlying biomechanics. Naturally, larger hammers should result in greater intersegmental moments for
primary degrees of freedom, e.g. shoulder elevation, elbow flexion, radial/ulnar deviation. However, related to the supposition on diversity of muscle forces, it is likely that moments for secondary degrees of freedom are also affected, e.g. shoulder rotation, elbow pronation, wrist flexion. As with lower-extremity, e.g. post anterior cruciate ligament ruptured knees (Noyes et al., 1992), these secondary moments are associated with effort and difficulty of control of the motion required to generate an efficient movement. Ultimately the secondary moments are generated or resisted by muscular forces or passive joint tissues, respectively (Winter 2009) and result in higher net loads applied to tissues responsible for maintaining joint stability. Both of these factors contribute to overall musculoskeletal loading and may have cumulative importance as related to fatigue or potential injury (Kumar, 2001). If increased moments are determined, further investigation into hammering could be warranted. Therefore, the purpose of this study was to use an existing opensource musculoskeletal model to determine intersegmental moments for wrist, elbow and shoulder using inverse-dynamics when subjects were engaged in a common hammering task with hammers of different mass. The hypotheses for this study were threefold. First, a larger mass hammer would result in a greater time duration for each phase of hammering and for the entirety of the motion. Second, a hammer with larger mass would result in increased intersegmental moments at wrist, elbow and shoulder. Third, a larger hammer would result in larger relative increases for moments associated with secondary degrees of freedom compared to relative increases in moments for primary degrees of freedom. 2. Methods 2.1. Subjects Ten male subjects (mean age: 24, range 22e29, height: mean 1.81 m, S.D. 0.04 m, mass: mean 80.7 kg, S.D. 12.1 kg) without history of dominant upper extremity neuro-musculoskeletal injury or surgery, and who had used a hammer on a least ten previous occasions to drive nails, were recruited from a university setting. All subjects were right hand dominant and granted informed written consent as required for ethics approval from the Institutional Review Board. 2.2. Procedure Following collection of height and mass, subjects were familiarized with the nature of the tasks and were required to practice the hammering tasks for a few hammer swings. The hammers were a standard 0.45 kg (1 pound) ball peen and a 0.91 kg (2 pound) sledge with handle lengths of 30 cm and 32.5 cm, respectively (Fig. 1). Nails were secured in a custom fixture affixed to the force plate. Position was standardized by instructing subjects to locate themselves facing the same global direction in a single knee on ground (kneeling) posture with the opposite (left) foot flat on the floor a comfortable distance away from the nail so that the nail could be reached with the hammer and driven as accurately and with as much force and power as possible while dedicating equal priority to each requirement such as a carpenter would drive a nail into a floor. Each nail was struck only once, and three nails were struck with each hammer. To mitigate any effects of learning or fatigue, the order of hammer usage was randomized across subjects. 2.3. Data collection Motion of markers attached to the upper extremity and the
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Fig. 1. Hammers investigated in this study were a common ball peen hammer (1 lb) and a sledge hammer (2 lb) (manufactured by Vaughan Manufacturing, Hebron, Illinois and Masterforce, Menards, Eau Claire, Wisconsin, respectively).
hammer was recorded during the hammer cocking, swinging and follow-through with a twelve camera optoelectronic system (Vicon Motion Capture) (100 Hz). Reflective markers were attached to: spinous process of C7, xiphoid process, proximal clavicle at sternoclavicular joint, bilateral acromion processes, bicep, medial and lateral humeral epicondyles, right forearm, radial and ulnar styloid
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processes, metacarpal phalangeal joints of middle (MCP3) and small (MCP5) fingers (Fig. 2). The bicep and forearm markers were applied to the midpoints of the anterior aspect of the upper and lower arm, respectively. One marker was attached to the hammer handle and two markers were attached to the hammer head. The three markers were used to locate hammer center of mass and measure hammer motion (Figs. 1 and 3). A static calibration trial was collected prior to hammering with subjects standing in a comfortable upright posture while holding the 1 lb hammer. The interaction of the hammer with the nail was recorded with a force platform (1000Hz) (AMTI-OR6-7-1000) (Fig. 2, top panels). Nails (16d) were held in a custom fixture rigidly clamped to the platform. Prior to subject experimentation, hammer masses and mass moments of inertia (MMOI) were measured (Table 1). Transverse (swingweight) and lateral (spinweight) MMOIs were determined by treating the hammer as a simple pendulum and the polar (twistweight) MMOI was determined using calibrated wire of known stiffness to suspend the hammer as a bifilar pendulum (Spurr et al., 2014). 2.4. Analysis Marker data were post-processed (Vicon Nexus) to ensure correct marker identification and that reaction forces and moments were measured by the platform. A dynamic musculoskeletal model of the upper extremity (Saul et al., 2014) capable of being utilized in forward- or inverse-dynamics analyses was scaled from the static calibration trial to calculate subject-specific joint centers and segment inertial properties from marker and anthropometric data
Fig. 2. Vicon Nexus model (top panels) and OpenSim (bottom panels) denoting transitions between phases of hammering motion. (A) Cocking phase began when subject initiated motion. (B) Swing phase begins instant after hammer reaches fully cocked position. (C) Strike/Follow-through phase begins when hammer contacts nail. Green vector (A & B, bottom panels) denotes force applied by hammer to hand (force and moment determined via inverse-dynamics, moment not shown). Red vectors (C) represent the force measured by plate (top) and applied by nail/ground to hammer (bottom). For C (similar to A & B) force and moment applied by hammer to hand was determined via inverse-dynamics (not shown).(For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)
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Fig. 3. Kinetic model of the hammer used in inverse-dynamics calculation to determine intersegmental force and moment (F HammerHand and M HammerHand , respectively) between hammer and hand. mHammer and aHammer are the mass and acceleration of the hammer; g is the acceleration due to gravity; H_ Hammer is the rate of change of hammer angular momentum; rMCP3 is a position vector of the MCP3 marker relative to hammer center of mass.
Table 1 Hammer inertial properties.
Mass [g] Swingweight [kg-cm2] Spinweight [kg-cm2] Twistweight [kg-cm2]
1 lb hammer
2 lb hammer
650 58 55 4
1216 124 117 10
using OpenSim (Delp et al., 2007). Each scaled upper extremity model's mass was calculated as a fixed fraction (6%) of subject's overall body mass (Chaffin and Andersson, 1999). The “preserve mass distribution” option within OpenSim scaled segment masses linearly with subject mass. Marker and force plate data were filtered with second order, 6 Hz, low-pass Butterworth filters. For each nail strike trial, joint angles describing shoulder, elbow and wrist motion were calculated with OpenSim's Inverse Kinematics. Three-dimensional hammer kinematics and accelerations were calculated in Matlab and were used in an inverse-dynamics Newton-Euler calculation (Vaughan et al., 1999) with hammer inertial properties to determine the intersegmental force and moment between the hammer and the hand (Fig. 3). By including the hammer force and moment (applied to the model hand at MCP3) and the reaction force and moment of the hammer with the nail measured by the force platform during strike phase and applied at the hammer head, the model generalized forces which reproduced the measured kinematics were calculated using OpenSim's Inverse Dynamics (Steele et al., 2014). The shoulder elevation angle, shoulder elevation and shoulder rotation generalized forces represent the corresponding moments at the shoulder, while the elbow flexion and pronation-supination generalized forces represent the corresponding moments at the elbow. For the wrist, the deviation and flexion generalized forces represent the corresponding joint moments. For the inversedynamics analysis, muscle forces and damping (ligament and joint coordinate limiting) present in the OpenSim model (Saul et al., 2014) were removed so that calculated moments only included the effects of gravity and accelerations on segment masses and inertia. By using the generalized forces as coordinate actuators, OpenSim's Joint Reaction analysis was used to calculate intersegmental reaction forces (Lerner et al., 2014). From the Joint Reaction analysis, the joint reaction loads (i.e. moment components) on the ulna in the forward ulnar direction and on the radius in the radial long axis
direction represented the elbow varus-valgus and wrist internalexternal rotation intersegmental moments, respectively. The nail strike motion was divided into three phases in order to consider temporal changes in the loadings (Fig. 2). The cocking phase began when the subject initiated motion and ended when the hammer reached the highest vertical point. The swing phase occurred from the instant after cocking phase ended until the force plate registered a 2 N vertical force. Strike/follow-through phase was from the instant after end of swing phase until motion ceased. Trial times and phase durations were calculated and analyzed with t-tests. For each phase, the maximum and minimum moments were identified and recorded. Repeated-measures ANOVAs (subject X hammer) were performed (Matlab, R2014a) to test for significant differences between hammers for each moment component at each joint. As such, subject was a blocking variable to account for intersubject differences.
3. Results Scaling of the generic OpenSim model resulted in subjectspecific models with unique anthropometric parameters (Table 2). Trial durations for 1 lb hammer (0.81 ± 0.11 s, mean ± S.D) were significantly shorter than for 2 lb hammer (0.99 ± 0.22 s) (p < 0.01). Likewise, durations were shorter for the smaller hammer for cocking phase (0.34 ± 0.07 s and 0.44 ± 0.18s) (p < 0.05), for swing phase (0.07 ± 0.04 s and 0.12 ± 0.07 s) and strike/follow-through (0.39 ± 0.03 s and 0.43 ± 0.03 s) (both p < 0.01). Intersegmental moments varied throughout the nail strike motion both within and between phases for the 1 lb (Fig. 4) and 2 lb (Fig. 5) hammers. The larger hammer generally resulted in increased peak intersegmental moments at all upper extremity
Table 2 Scaled model segment masses. Mass [kg] (S.D.) Upper-extremity Clavicle Scapula Humerus Ulna Radius Hand
4.53 0.15 0.67 1.90 1.05 0.22 0.55
(0.89) (0.03) (0.13) (0.37) (0.21) (0.04) (0.11)
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Fig. 4. Intersegmental moments for (A) shoulder, (B) elbow and (C) wrist for 1 lb hammer. Grey denotes individual trials; black represents ±1 standard deviation from mean. Plane of elevation moment is positive counter-clockwise on right shoulder; other moments are positive for labeled direction. Moments are piece-wise linear length normalized to the average trial and phase durations for 1 lb hammers.
joints although the increase differed in terms of relative magnitude and significance of the effect (Fig. 6). Absolute measured moments and changes due to the heavier hammer were largest at the shoulder, followed by the elbow, and then the wrist. The largest moments were for shoulder plane of elevation and were somewhat smaller for shoulder elevation, while elbow flexion moments were substantially larger than other elbow moments for all phases of hammering. At the wrist during cocking and swinging phases, the largest moments were for ulnar deviation, while during strike/ follow-through, radial deviation tended to be relatively large with other moments also considerably increased. In general, increases in peak moments tended to be more pronounced in cocking and striking/follow-through phases as opposed to swinging phase. There was a trend toward larger relative increases in moments for elbow and wrist as compared to shoulder: 119% and 99% compared to 68% on average, respectively (Fig. 6). More specifically, at the shoulder the heavier hammer resulted in a significantly increased maximum elevation moment (by 9%) during the cocking (29.9 ± 2.5 vs. 27.5 ± 2.9 N-m, mean ± standard
error, p ¼ 0.0001) and strike/follow-through (89%) phases (55.8 ± 13.6 vs. 29.4 ± 2.8 N-m, p ¼ 0.002), while for internal rotation a significant increase (1.4 ± 0.1 vs. 0.9 ± 0.1 N-m, p ¼ 0.01) was found (53%) during the cocking phase. Likewise, the minimum moments for these joint angles (representing shoulder depression and external rotation, respectively) were also significantly different; the larger hammer resulted in increased shoulder depression moment during the strike phase (71.9 ± 10.6 vs. 49.7 ± 2.8 N-m, p ¼ 0.004) and increased external rotation moment (192%) in cocking phase (1.1 ± 0.1 vs. 0.4 ± 0.1 N-m, p < 0.0001). No significant differences were detected in peak shoulder elevation angle moments. At the elbow, the changes in peak moments were not as large in magnitude, but generally were more consistent across joint angles and demonstrated greater relative changes (Fig. 6B). With the heavier hammer, maximum elbow flexion moments across all phases were significantly greater (all p < 0.0001) by an average of 77% (e.g. for cocking, 14.5 ± 0.2 vs. 8.4 ± 0.1 N-m), while minimum elbow flexion moments were also significantly different with
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Fig. 5. Intersegmental moments for (A) shoulder, (B) elbow and (C) wrist for 2 lb hammer. Grey denotes individual trials; black represents ±1 standard deviation from mean. Plane of elevation moment is positive counter-clockwise on right shoulder; other moments are positive for labeled direction. Moments are piece-wise linear length normalized to the average trial and phase durations for 2 lb hammers.
mixed effect (p ¼ 0.03, p < 0.0001, p < 0.001, for phases respectively). For the cocking and swing phases, the peak minimum elbow flexion moments were in flexion direction, while for the strike/ follow-through phase the minimum moment was in extension. The peak maximum elbow pronation moment was significantly increased (average 120%) during cocking (0.5 ± 0.0 vs. 0.2 ± 0.0 Nm, p < 0.001) and strike/follow-through (2.4 ± 0.4 vs. 1.1 ± 0.1 N-m, p ¼ 0.03). Likewise, the peak minimum elbow pronation moment (i.e. supination) was significantly increased for all three phases by an average 143%) (1.7 ± 0.2 vs. 0.8 ± 0.1 N-m for strike/followthrough, p ¼ 0.001). The larger hammer caused a significant increase (138%) in elbow varus moment for the cocking phase (0.8 ± 0.1 vs. 0.3 ± 0.1 N-m, p ¼ 0.015) while the peak minimum varus (i.e. valgus) moment was significantly increased for all phases by an average 144% (e.g. for strike/follow-through, 7.4 ± 1.4 vs. 3.1 ± 0.3 N-m, p ¼ 0.04). When compared to the moments and changes at the shoulder and elbow, wrist peak moments and changes were generally small in magnitude, but for some components still demonstrated rather
large effect (Fig. 6C). The heavier hammer caused increased (by 99%) wrist flexion moment for strike/follow-through (1.9 ± 0.2 vs. 1.0 ± 0.1 N-m, p ¼ 0.0001), whereas the peak minimum moment (representing wrist extension) was significantly increased for all phases on average 83% (for strike/follow-through, 2.4 ± 0.2 vs. 1.3 ± 0.1 N-m, p < 0.001). Peak maximum ulnar deviation moment was significantly increased with the larger hammer during the swing (1.6 ± 0.2 vs. 0.6 ± 0.0 N-m, p < 0.001) and strike/followthrough phases (5.3 ± 0.5 vs. 2.7 ± 0.1 N-m, p < 0.0001). The peak minimum ulnar moments were in radial direction for all phases and were significantly increased (by average 108%) with 2 lb hammer (cocking phase: 2.5 ± 0.3 vs. 1.2 ± 0.1 N-m, and p < 0.001). The measured peak maximum wrist rotation moments were in the internal direction for the cocking and strike/follow-through phases and were significantly increased on average 226% with the larger hammer with large effect (for strike/follow-through, 2.3 ± 0.2 vs. 1.0 ± 0.1 N-m, p < 0.0001). Lastly, peak minimum wrist rotation moments were all negative (representing external rotation) for all phases and were significantly increased (74% on average) with the
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Fig. 6. Mean (±standard error) peak (A) shoulder, (B) elbow and (C) wrist intersegmental moments [N-m] for 1 and 2 pound hammers during phases of driving a nail. Significant differences denoted with * for p < 0.05 and ** for p < 0.01. Negative shoulder moments represent: elevation plane in counter-clockwise direction, depression, external rotation. Negative elbow moments represent: valgus, extension, supination. Negative wrist moments represent: radial deviation, extension, external rotation. (Note difference in scale for shoulder internal rotation and elbow varus and pronation moments).
large hammer (for strike/follow-through, 2.8 ± 0.3 vs. 1.4 ± 0.1 N-m, p ¼ 0.005). 4. Discussion This study investigated effects of increased hammer mass and moment of inertia on upper extremity joint kinetics during hammering of a nail. The motivation for this study is the hypothesis that larger hammers cause greater loading to the upper extremity musculoskeletal tissues. Results determined here that the larger hammer resulted in increased intersegmental joint moments for some components and some phases of the hammering motion support this hypothesis. Relatively large changes in intersegmental moments (50 to > 150%) were recorded when hammer moments of inertia were increased in the range of 113e150%. Additionally, trial time and phase durations were significantly greater (~10e30%) and greater variability (~50%) was demonstrated in these times for the larger hammer. Moments were generally smaller in absolute magnitude than reported in previous studies of upper extremity (Creveaux et al., 2013a,b; Rogowski et al., 2014; Slavens et al., 2010) but were similar to these studies and those of other joints, e.g. (Lerner et al., 2014), in that larger moments were calculated for primary degrees of freedom associated with generating the motion.
In addition, the hypothesis that a larger hammer would result in greater relative increases in moments for secondary degrees of freedom was supported. Taken together, the changes in moments and timing indicate greater loads on joints and musculoskeletal tissues for an extended period with the larger hammer. Subjects likely found the larger hammer more challenging to control as they required more time on average and demonstrated greater variability in duration for cocking and swinging phases than for the smaller hammer. Differences in duration might also have been related to novice subjects engaged in hammering. By definition, novices generate greater variance than more experienced subjects e.g. two subjects tended to generate larger moments in relation to others (Figs. 4 and 5). Subjects with less expertise likely experienced greater difficulty executing the coordinated motion across multiple joints required to generate large moments which resulted in the few differences found for shoulder moments, e.g. no differences were found at the shoulder during swing phase as subjects might have tended to let the hammer fall rather than actively extending the shoulder. In addition, the task was not representative of all hammering. This strategy and other potential differences described here could be ascertained in the future by measuring muscle activity during hammering. By assigning a task of striking a small nail and thus
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requiring novice subjects to optimize accuracy as well as force/ power, the resulting motion was of low acceleration when compared to a task where accuracy is not as paramount, such as striking a larger target, e.g. driving a stake of large cross-section where greater acceleration and force delivery to the target would be achieved. Such a task would likely result in greater loads imparted to the upper extremity for all phases, particularly when using a larger hammer, and might result in greater differences between the hammers. It is important to realize that for many of the secondary degrees of freedom where large relative changes in moments were found for some joints and components, the absolute changes in joint moments were small during cocking and swing phases, making the clinical significance of these results less likely and more difficult to assess. Even though the absolute increases were small, they were statistically significant for many components and there was a distinct trend towards increased moments when using the larger hammer that might be more pronounced for a different, more representative hammering task as mentioned above. Absolute increases were however larger for most joints and components during strike/follow through, indicating greater potential for injury during this phase. The relatively large increase for shoulder elevation during strike/follow-through indicates potential for higher concentric load placed on rotator cuff muscles and tendons, particularly supraspinatus, when hammer strikes a target. The increased depression moment during strike follow-through likely causes eccentric loading on rotator cuff, and to some extent deltoid, as these muscles experience lengthening contractions while slowing or arresting the hammer. Additionally, long head of biceps and its tendon would experience increased loading as it is partially responsible for aiding in humeral head depression while primarily responsible for the increased elbow flexion observed in all phases. Future studies to measure EMG of these and other muscles during hammering could clarify these potential injury mechanisms as previously performed for tennis racket swing (Giangarra et al., 1993). At the elbow during strike/follow through increased varusvalgus moments indicate greater tensile loads applied to radial and ulnar collateral ligaments, respectively, and also compressive loads to articular cartilage due to resulting force imbalance and inability of ligaments to fully resist the loading (Morrey and An, 1983). Increased elbow flexion moment results in greater loads applied to biceps tendon in all phases and especially with increased extension moment during strike/follow through when biceps could experience an eccentric contraction. At elbow, increased pronation and supination moments in strike/follow through require increased forces from pronators teres and quadratus and forearm supinator and biceps, respectively, thus increasing the potential for injuries to these muscles especially with repetitive use. For the wrist, increased radial and ulnar deviation moments during strike phase combined with greater flexion and extension moments would overload flexor carpi ulnaris and radialis as well as extensors carpi radialis longus, radialis brevis and ulnaris as commonly observed in clinical settings. The conjoined tendons for these muscles likely experience disorders in a manner similar to golf and tennis overuse injuries, i.e. radial and ulnar epicondylitis, due to eccentric contractions experienced when subjected to relatively large moments of changing direction during the short duration strike/follow through phase (Creveaux et al., 2013a,b; Knudson and Bahamonde, 1997). The trend of larger relative increases in moments for elbow and wrist when using greater mass hammers supports and agrees with the limited but similar findings of increased activity for forearm muscles as compared to upper arm and shoulder muscles when workers are required to use tools having more degrees of freedom which provide greater mechanical, physiological and psychological challenge (Fischer et al., 2009). However, in
comparing across joints, it is important to remember these increases were larger in a relative sense; absolute increases were larger for proximal joints. Because the hammers used in this study were of different mass, the overall magnitude of inertia differed accordingly, but remarkably the symmetries of inertia were nearly identical. The magnitude and symmetry of inertia for objects moved in space have been assumed to affect the intensity and patterning of muscle forces, respectively (Fitzpatrick et al., 2012). It would be interesting to investigate whether hammers with more varying magnitudes and particularly symmetries of inertia would require different motor control, and thus affect loads on upper extremity joints resulting in greater differences in kinematics and accelerations for both primary and especially secondary degrees of freedom. Additional work is necessary to understand hammering in these more representative, and likely higher loading situations. There are several limitations of this work related to the model and experimental task. First, although a generic model of the upper extremity was scaled to anthropometry of subjects, little information is available in literature to document the accuracy of this process. For lower-extremity, intersegmental forces and moments are usually normalized by body weight and product of body weight and height, respectively. Given the substantial inter-subject variability (Figs. 4 and 5) and the absence of similar scaling laws for upper-extremity, non-normalized moments were analysed statistically by blocking across subjects in order to account for intersubject variability. Next, although the interaction between the hammer and hand occurs as a distributed contact stress, by modeling all segments as rigid bodies as required for an inverse dynamics analysis, the force and moment between hammer and hand were assumed to be concentrated at a point on the hand and therefore calculated directly from hammer inertial properties and kinematics. In order to determine contact stresses between hammer and hand and at other joints, a deformable analysis, e.g. finite element analysis, would be required. Lastly, differences in handle size and shape were not considered here but have been shown to affect wrist and hand muscle forces and related load sharing (Rossi et al., 2012, Rossi et al., 2015). In conclusion, previous studies have suggested that mechanical loading is a risk factor for upper extremity injury (Buckle and Devereux, 2002). This study found increased intersegmental moments when using a larger hammer. Increases in joint moments when using a larger hammer were small in absolute magnitude for some joints, but there was a trend towards larger relative moments for the elbow and wrist, and particularly for secondary degrees of freedom such as elbow varus-valgus and pronation-supination and wrist rotation and radial-ulnar deviation. With the larger hammer, subjects required more time for the hammering motion and more time within phases with greater variability of the timing. These data and findings support previous understanding of the deleterious effects of hammer usage on upper extremity musculoskeletal tissues, and provide measurements of potential loads and describe mechanisms of the loadings. This information can potentially help explain an increased risk of developing cumulative trauma disorders due to repetitive hammer use and be used to inform future studies on hammering. Conflict of interest statement No author of this work has any financial or other conflict of interest. Acknowledgements This study was made possible and supported by equipment
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purchased under NSF-CBE-MRI 1337511. Sponsors had no role in: study design, data collection or analysis, interpretation of data, in manuscript writing or decision to submit the manuscript for publication. The authors thank Katherine Saul and Meghan Vidt for assistance with implementation of the MoBL-ARMs model. References Bernstein, N., 1967. The Coordination and Regulation of Movement. Pergamon Press, London. Buckle, P.W., Devereux, J.J., 2002. The nature of work-related neck and upper limb musculoskeletal disorders. Appl. Ergon. 33, 207e217. Chaffin, D.B., Andersson, G.B.J., 1999. Anthropometry in Occupational Biomechanics. In: Occupational Biomechanics. John Wiley and Sons, New York, NY. Cote, J.N., Feldman, A.G., Mathieu, P.A., Levin, M.F., 2008. Effects of fatigue on intermuscular coordination during repetitive hammering. Mot. Control 12, 79e92. Cote, J.N., Raymond, D., Mathieu, P.A., Feldman, A.G., Levin, M.F., 2005. Differences in multi-joint kinematic patterns of repetitive hammering in healthy, fatigued and shoulder-injured subjects. Clin. Biomech. 20, 581e590. , P., Che ze, L., Rogowski, I., 2013a. Joint Creveaux, T., Dumas, R., Hautier, C., Mace kinetics to assess the influence of the racket on a tennis player's shoulder. J. Sports Sci. Med. 12, 259e266. , P., Che ze, L., Rogowski, I., 2013b. Influence of racket Creveaux, T., Dumas, R., Mace polar moment on joint loads during tennis forehand drive. Comput. Methods Biomechancis Biomed. Eng. 16, 99e101. Delp, S.L., Anderson, F.C., Arnold, A.S., Loan, P., Habib, A., John, C.T., Guendelman, E., Thelen, D.G., 2007. OpenSim: open-source software to create and analyze dynamic simulations of movement. IEEE Trans. Biomed. Eng. 55, 1940e1950. Fischer, S.L., Wells, R.P., Dickerson, C.R., 2009. The effect of added degrees of freedom and handle type on upper limb muscle activity during simulated hand tool use. Ergonomics 52, 25e35. Fitzpatrick, P., Wagman, J.B., Schmidt, R.C., 2012. Alterations in movement dynamics in a tool-use task: the role of action-relevant inertial tool parameters. Z. fur Psychol. 220, 23e28. Giangarra, C.E., Conroy, B., Jobe, F.W., Pink, M., Perry, J., 1993. Electromyographic and cinematographic analysis of elbow function in tennis players using single- and double-handed backhand strokes. Am. J. Sports Med. 21, 394e399. Hammarskjold, E., Harms-Ringdahl, K., 1992. Effect of arm-shoulder fatigue on carpenters at work. Eur. J. Appl. Physiol. Occup. Physiol. 64, 402e409. Kahrs, B.A., Jung, W.P., Lockman, J.J., 2014. When does tool use become distinctively human? Hammering in young children. Child. Dev. 85, 1050e1061. Karkowski, W., Mattila, M., Quesada, P.M., Kwon, Y.G., James, D., 2002/2003. The effects of hammer design and hammering surface orientation on task performance. Occup. Ergon. 3, 45e54. Khan, A.A., Shakeb, M., 2009. Effect of posture, length of hammer-handle and frequency on discomfort. In: Developments in Agricultural and Industrial Ergonomics: General Studies, Vol. 1. Allied Publishers, Delhi. Knowlton, R.G., Gilbert, J.C., 1983. Ulnar deviation and short-term strength reductions as affected by a curve-handled ripping hammer and conventional claw hammer. Ergonomics 26, 171e179. Knudson, D., Bahamonde, R.E., 1997. Upper extremity angular kinematics of the one-handed backhand drive in tennis players with and without tennis elbow. Int. J. Sports Med. 18, 79e82.
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Konz, S., 1986. Bent Hammer Handles. Hum. Factors J. Hum. Factors Ergon. Soc. 28, 317e323. Kumar, S., 2001. Theories of musculoskeletal injury causation. Ergonomics 44, 17e47. Lerner, Z.F., Haight, D.J., DeMers, M.S., Board, W.J., Browning, R.C., 2014. The effects of walking speed on tibiofemoral loading estimated via musculoskeletal modeling. J. Appl. Biomechanics 30, 197e205. Leventhal, E.L., Moore, D.C., Akelman, E., Wolfe, S.W., Crisco, J.J., 2010. Carpal and forearm kinematics during a simulated hammering task. J. Hand Surg. Am 35, 1097e1104. Mattila, M., Karkowski, W., Vilkki, M., 1993. Analysis of working postures in hammering tasks on building construction sites using the computerized OWAS method. Appl. Ergon. 24, 405e412. Morrey, B., An, K.N., 1983. Articular and ligamentous contributions to the stability of the elbow joint. Am. J. Sports Med 11, 315e319. Noyes, F.R., Schipplein, O.D., Andriacchi, T.P., Saddemi, S.R., Weise, M., 1992. The anterior cruciate ligament-deficient knee with varus alignment. An analysis of gait adaptations and dynamic joint loadings. Am. J. Sports Med. 20, 707e716. Putnam, C.A., Jenkins, S., 1993. Effect of Hammer Load on Segment Motion Patterns and Segment Interaction in a Hammering Task. XIV Congress of the International Society of Biomechanics Paris, France. Rogowski, I., Creveaux, T., Cheze, L., Mace, P., Dumas, R., 2014. Effects of the racket polar moment of inertia on dominant upper limb joint moments during tennis serve. PLoS One 9, e104785. lot, L., Barla, C., Vigouroux, L., 2012. Characterisation of forces Rossi, J., Berton, E., Gre exerted by the entire hand during the power grip: effect of the handle diameter. Ergonomics 55, 682e692. Rossi, J., Goislard De Monsabert, B., Berton, E., Vigouroux, L., 2015. Handle shape effects the grip force distribution and the muscle loadings during power grip tasks. J. Appl. Biomechanics 31, 430e438. Saul, K.R., Hu, X., Goehler, C.M., Vidt, M.E., Daly, M., Velisar, A., Murray, W.M., 2014. Benchmarking of dynamic simulation predictions in two software platforms using an upper limb musculoskeletal model. Comput. Methods Biomechancis Biomed. Eng. 18, 1445e1458. Schoenmarklin, R.W., Marras, W.S., 1989a. Effects of handle angle and work orientation on hammering: I Wrist motion and hammering performance. Hum. Factors J. Hum. Factors Ergon. Soc. 31, 397e411. Schoenmarklin, R.W., Marras, W.S., 1989b. Effects of handle angle and work orientation on hammering: II Muscle fatigue and subjective ratings of body discomfort. Hum. Factors J. Hum. Factors Ergon. Soc. 31, 413e420. Slavens, B.A., Sturm, P.F., Harris, G.F., 2010. Upper extremity inverse dynamics model for crutch-assisted gait assessment. J. Biomechanics 43, 2026e2031. Spielholz, P., Wiker, S.F., Silverstein, B., 1998. An ergonomic characterization of work in concrete form construction. Am. Industrial Hyg. Assoc. J. 59, 629e635. Spurr, J., Goodwill, S., Kelley, J., Haake, S., 2014. Measuring the inertial properties of a tennis racket. Procedia Eng. 72, 569e574. Steele, K.M., DeMers, M.S., Schwartz, M.H., Delp, S.L., 2014. Compressive tibiofemoral force during crouch gait. Gait Posture 35, 556e560. Vaughan, C.L., Davis, B.L., O'Connor, J.C., 1999. Integration of Anthropometry, Displacements, and Ground Reaction Forces. In: Dynamics of Human Gait. Human Kinetics Publishers, Champaign, IL, pp. 15e131. Wagman, J.B., Carello, C., 2003. Haptically creating affordances: the user-tool interface. J. Exp. Psychol. Appl. 9, 175e186. Winter, D.A., 2009. Biomechanics and Motor Control of Human Movement. John Wiley & Sons, New York, NY.
Contents lists available at ScienceDirect
Applied Ergonomics journal homepage: www.elsevier.com/locate/apergo
Effect of hammer mass on upper extremity joint moments Nilanthy Balendra a, Joseph E. Langenderfer b, * a b
Department of Physical Therapy, Central Michigan University, Mount Pleasant, MI 48859, USA School of Engineering and Technology, Central Michigan University, Mount Pleasant, MI 48859, USA
a r t i c l e i n f o
a b s t r a c t
Article history: Received 3 February 2016 Received in revised form 29 November 2016 Accepted 2 December 2016
This study used an OpenSim inverse-dynamics musculoskeletal model scaled to subject-specific anthropometrics to calculate three-dimensional intersegmental moments at the shoulder, elbow and wrist while 10 subjects used 1 and 2 lb hammers to drive nails. Motion data were collected via an optoelectronic system and the interaction of the hammer with nails was recorded with a force plate. The larger hammer caused substantial increases (50e150%) in moments, although increases differed by joint, anatomical component, and significance of the effect. Moment increases were greater in cocking and strike/follow-through phases as opposed to swinging and may indicate greater potential for injury. Compared to shoulder, absolute increases in peak moments were smaller for elbow and wrist, but there was a trend toward larger relative increases for distal joints. Shoulder rotation, elbow varus-valgus and pronation-supination, and wrist radial-ulnar deviation and rotation demonstrated large relative moment increases. Trial and phase durations were greater for the larger hammer. Changes in moments and timing indicate greater loads on musculoskeletal tissues for an extended period with the larger hammer. Additionally, greater variability in timing with the larger hammer, particularly for cocking phase, suggests differences in control of the motion. Increased relative moments for distal joints may be particularly important for understanding disorders of the elbow and wrist associated with hammer use. © 2016 Published by Elsevier Ltd.
Keywords: Inverse-dynamics Kinetics
1. Introduction Throughout human history hammering has been a common manual task afforded by holding an object of relatively large concentrated mass in, or, by use of a handle, at some extended distance from the hand(s) (Kahrs et al., 2014). As the mass of the object is increased, and/or the object is moved further from the hand, the mass moment of inertia (MMOI) increases and the potential for the human to exert impulsive forces on the external environment and thereby perform work is consequently improved. However, the potential cost of this increased MMOI and associated work may be increased loads on the human body in the form of joint loadings and stresses applied to musculoskeletal tissues. In contrast to medieval times when a blacksmith formed metal into useful objects via repetitive hammering, modern industrial processes and the people that perform them have benefited from electromechanical automation of work. However, hammering remains an important task in carpentry and concrete form
* Corresponding author. E-mail address: [email protected] (J.E. Langenderfer). http://dx.doi.org/10.1016/j.apergo.2016.12.001 0003-6870/© 2016 Published by Elsevier Ltd.
construction, where it is the single most commonly performed work at 17% of the workday (Spielholz et al., 1998). Additionally, workers rate arm strain to be either high or very high in 52% of hammering tasks (Mattila et al., 1993). While the nature of hammering in terms of intensity and body position tends to be variable rather than fixed (Spielholz et al., 1998), by definition, hammer usage is extremely repetitive (Bernstein, 1967) and frequently occurs in a compromised overall posture which tends to cause cumulative fatigue related injuries (Mattila et al., 1993; Spielholz et al., 1998). Considerable attention has been devoted to investigating hammering from several perspectives. Alterations in hammer design in terms of mass and handle shape or length have been shown to affect many outcomes, including: ability to perform and complete hammering tasks (Konz, 1986; Schoenmarklin and Marras, 1989a,b), perceived upper extremity exertion (Karkowski et al., 2002/2003), arm muscle fatigue (Schoenmarklin and Marras, 1989a,b), body discomfort (Khan and Shakeb, 2009; Schoenmarklin and Marras, 1989a,b), and changes in grip strength before and after hammering (Knowlton and Gilbert, 1983). Additionally, when the orientation of the work surface is horizontal, rather than vertical, the human is able to attain better
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performance of the hammering task (Karkowski et al., 2002/2003, Schoenmarklin and Marras, 1989a,b) with decreased muscle exertion and discomfort (Schoenmarklin and Marras, 1989a,b). Previous investigations have observed wrist (Schoenmarklin and Marras, 1989a,b) and upper extremity (Cote et al., 2008, Cote et al., 2005) kinematics during actual (Cote et al., 2008, Cote et al., 2005; Schoenmarklin and Marras, 1989a,b) or simulated (Leventhal et al., 2010) hammering using electromechanical (Schoenmarklin and Marras, 1989a,b), optoelectronic (Cote et al., 2008, Cote et al., 2005) or imaging techniques (Leventhal et al., 2010). In general, angled hammer handles have been shown to reduce ulnar deviation during swinging phase (Knowlton and Gilbert, 1983; Schoenmarklin and Marras, 1989a,b), but increase radial deviation at the initial hammering position (Schoenmarklin and Marras, 1989a,b). When compared with other tasks, nailing with a hammer has been perceived to be more difficult following fatigue (Hammarskjold and Harms-Ringdahl, 1992). Conversely, when repetitive hammering has been used as a fatiguing activity, it has been shown to cause reduced elbow velocity, acceleration, range of motion and grip strength (Cote et al., 2008, Cote et al., 2005) while leaving wrist and shoulder kinematics and upper extremity EMG unaffected. When compared to hammering, tennis is not as common an activity among the general population in terms of frequency and duration of performance, but is similar in that a tool is used as an extension of the upper extremity in order to accomplish a manual task. Tennis racket mass and MMOI play an important role in the ability of players to control the racket during swing, ball strike and follow through and thus impart energy to a ball in order to return it to an opponent. Relatively small changes (10e15%) in racket mass and MMOI have been shown to substantially affect joint loads with calculated differences of 30e40% (Creveaux et al., 2013a,b; Rogowski et al., 2014). Hammers used in common carpentry tasks (455e900 g) generally have greater mass (100e200%) but smaller MMOIs (from 25 to 50%, measured in this study) compared to tennis rackets (Creveaux et al., 2013a,b). If one accepts that hammering is a relatively common activity among professional laborers and the general population, the question of upper extremity joint loading resulting from hammer usage becomes relevant. In spite of the ubiquitous nature of hammer use among lay persons and professional manual laborers, and the substantial number of previous investigations of hammering, estimates of the loads applied to upper extremity joints resulting from hammer tool usage are notably absent from scientific literature. Effects of hammer mass and MMOI on timing, joint kinematics and moments have been examined previously with only very brief and undetailed results (Putnam and Jenkins, 1993). In ecological psychology, previous studies of object manipulation have explored the relationship between intended object tool usage and resultant grasping location (Fitzpatrick et al., 2012; Wagman and Carello, 2003). Subjects asked to grasp objects in a manner preparatory to hammering do so in a way that minimizes the symmetry of object inertia relative to magnitude of overall inertia. As such, subjects are described as behaving in a way that minimizes diversity of muscle forces required to control motion relative to the amount of force required to move the object in order to generate an efficient hammering motion. However, as these studies explored human behavior and not motor control or biomechanics, any hypothesis on muscle forces is unsupported as little is known about how actual hammering affects humans in terms of timing, kinematics, or joint kinetics. An investigation to determine how different hammers affect upper extremity kinetics is a necessary initial step towards understanding the underlying biomechanics. Naturally, larger hammers should result in greater intersegmental moments for
primary degrees of freedom, e.g. shoulder elevation, elbow flexion, radial/ulnar deviation. However, related to the supposition on diversity of muscle forces, it is likely that moments for secondary degrees of freedom are also affected, e.g. shoulder rotation, elbow pronation, wrist flexion. As with lower-extremity, e.g. post anterior cruciate ligament ruptured knees (Noyes et al., 1992), these secondary moments are associated with effort and difficulty of control of the motion required to generate an efficient movement. Ultimately the secondary moments are generated or resisted by muscular forces or passive joint tissues, respectively (Winter 2009) and result in higher net loads applied to tissues responsible for maintaining joint stability. Both of these factors contribute to overall musculoskeletal loading and may have cumulative importance as related to fatigue or potential injury (Kumar, 2001). If increased moments are determined, further investigation into hammering could be warranted. Therefore, the purpose of this study was to use an existing opensource musculoskeletal model to determine intersegmental moments for wrist, elbow and shoulder using inverse-dynamics when subjects were engaged in a common hammering task with hammers of different mass. The hypotheses for this study were threefold. First, a larger mass hammer would result in a greater time duration for each phase of hammering and for the entirety of the motion. Second, a hammer with larger mass would result in increased intersegmental moments at wrist, elbow and shoulder. Third, a larger hammer would result in larger relative increases for moments associated with secondary degrees of freedom compared to relative increases in moments for primary degrees of freedom. 2. Methods 2.1. Subjects Ten male subjects (mean age: 24, range 22e29, height: mean 1.81 m, S.D. 0.04 m, mass: mean 80.7 kg, S.D. 12.1 kg) without history of dominant upper extremity neuro-musculoskeletal injury or surgery, and who had used a hammer on a least ten previous occasions to drive nails, were recruited from a university setting. All subjects were right hand dominant and granted informed written consent as required for ethics approval from the Institutional Review Board. 2.2. Procedure Following collection of height and mass, subjects were familiarized with the nature of the tasks and were required to practice the hammering tasks for a few hammer swings. The hammers were a standard 0.45 kg (1 pound) ball peen and a 0.91 kg (2 pound) sledge with handle lengths of 30 cm and 32.5 cm, respectively (Fig. 1). Nails were secured in a custom fixture affixed to the force plate. Position was standardized by instructing subjects to locate themselves facing the same global direction in a single knee on ground (kneeling) posture with the opposite (left) foot flat on the floor a comfortable distance away from the nail so that the nail could be reached with the hammer and driven as accurately and with as much force and power as possible while dedicating equal priority to each requirement such as a carpenter would drive a nail into a floor. Each nail was struck only once, and three nails were struck with each hammer. To mitigate any effects of learning or fatigue, the order of hammer usage was randomized across subjects. 2.3. Data collection Motion of markers attached to the upper extremity and the
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Fig. 1. Hammers investigated in this study were a common ball peen hammer (1 lb) and a sledge hammer (2 lb) (manufactured by Vaughan Manufacturing, Hebron, Illinois and Masterforce, Menards, Eau Claire, Wisconsin, respectively).
hammer was recorded during the hammer cocking, swinging and follow-through with a twelve camera optoelectronic system (Vicon Motion Capture) (100 Hz). Reflective markers were attached to: spinous process of C7, xiphoid process, proximal clavicle at sternoclavicular joint, bilateral acromion processes, bicep, medial and lateral humeral epicondyles, right forearm, radial and ulnar styloid
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processes, metacarpal phalangeal joints of middle (MCP3) and small (MCP5) fingers (Fig. 2). The bicep and forearm markers were applied to the midpoints of the anterior aspect of the upper and lower arm, respectively. One marker was attached to the hammer handle and two markers were attached to the hammer head. The three markers were used to locate hammer center of mass and measure hammer motion (Figs. 1 and 3). A static calibration trial was collected prior to hammering with subjects standing in a comfortable upright posture while holding the 1 lb hammer. The interaction of the hammer with the nail was recorded with a force platform (1000Hz) (AMTI-OR6-7-1000) (Fig. 2, top panels). Nails (16d) were held in a custom fixture rigidly clamped to the platform. Prior to subject experimentation, hammer masses and mass moments of inertia (MMOI) were measured (Table 1). Transverse (swingweight) and lateral (spinweight) MMOIs were determined by treating the hammer as a simple pendulum and the polar (twistweight) MMOI was determined using calibrated wire of known stiffness to suspend the hammer as a bifilar pendulum (Spurr et al., 2014). 2.4. Analysis Marker data were post-processed (Vicon Nexus) to ensure correct marker identification and that reaction forces and moments were measured by the platform. A dynamic musculoskeletal model of the upper extremity (Saul et al., 2014) capable of being utilized in forward- or inverse-dynamics analyses was scaled from the static calibration trial to calculate subject-specific joint centers and segment inertial properties from marker and anthropometric data
Fig. 2. Vicon Nexus model (top panels) and OpenSim (bottom panels) denoting transitions between phases of hammering motion. (A) Cocking phase began when subject initiated motion. (B) Swing phase begins instant after hammer reaches fully cocked position. (C) Strike/Follow-through phase begins when hammer contacts nail. Green vector (A & B, bottom panels) denotes force applied by hammer to hand (force and moment determined via inverse-dynamics, moment not shown). Red vectors (C) represent the force measured by plate (top) and applied by nail/ground to hammer (bottom). For C (similar to A & B) force and moment applied by hammer to hand was determined via inverse-dynamics (not shown).(For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)
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Fig. 3. Kinetic model of the hammer used in inverse-dynamics calculation to determine intersegmental force and moment (F HammerHand and M HammerHand , respectively) between hammer and hand. mHammer and aHammer are the mass and acceleration of the hammer; g is the acceleration due to gravity; H_ Hammer is the rate of change of hammer angular momentum; rMCP3 is a position vector of the MCP3 marker relative to hammer center of mass.
Table 1 Hammer inertial properties.
Mass [g] Swingweight [kg-cm2] Spinweight [kg-cm2] Twistweight [kg-cm2]
1 lb hammer
2 lb hammer
650 58 55 4
1216 124 117 10
using OpenSim (Delp et al., 2007). Each scaled upper extremity model's mass was calculated as a fixed fraction (6%) of subject's overall body mass (Chaffin and Andersson, 1999). The “preserve mass distribution” option within OpenSim scaled segment masses linearly with subject mass. Marker and force plate data were filtered with second order, 6 Hz, low-pass Butterworth filters. For each nail strike trial, joint angles describing shoulder, elbow and wrist motion were calculated with OpenSim's Inverse Kinematics. Three-dimensional hammer kinematics and accelerations were calculated in Matlab and were used in an inverse-dynamics Newton-Euler calculation (Vaughan et al., 1999) with hammer inertial properties to determine the intersegmental force and moment between the hammer and the hand (Fig. 3). By including the hammer force and moment (applied to the model hand at MCP3) and the reaction force and moment of the hammer with the nail measured by the force platform during strike phase and applied at the hammer head, the model generalized forces which reproduced the measured kinematics were calculated using OpenSim's Inverse Dynamics (Steele et al., 2014). The shoulder elevation angle, shoulder elevation and shoulder rotation generalized forces represent the corresponding moments at the shoulder, while the elbow flexion and pronation-supination generalized forces represent the corresponding moments at the elbow. For the wrist, the deviation and flexion generalized forces represent the corresponding joint moments. For the inversedynamics analysis, muscle forces and damping (ligament and joint coordinate limiting) present in the OpenSim model (Saul et al., 2014) were removed so that calculated moments only included the effects of gravity and accelerations on segment masses and inertia. By using the generalized forces as coordinate actuators, OpenSim's Joint Reaction analysis was used to calculate intersegmental reaction forces (Lerner et al., 2014). From the Joint Reaction analysis, the joint reaction loads (i.e. moment components) on the ulna in the forward ulnar direction and on the radius in the radial long axis
direction represented the elbow varus-valgus and wrist internalexternal rotation intersegmental moments, respectively. The nail strike motion was divided into three phases in order to consider temporal changes in the loadings (Fig. 2). The cocking phase began when the subject initiated motion and ended when the hammer reached the highest vertical point. The swing phase occurred from the instant after cocking phase ended until the force plate registered a 2 N vertical force. Strike/follow-through phase was from the instant after end of swing phase until motion ceased. Trial times and phase durations were calculated and analyzed with t-tests. For each phase, the maximum and minimum moments were identified and recorded. Repeated-measures ANOVAs (subject X hammer) were performed (Matlab, R2014a) to test for significant differences between hammers for each moment component at each joint. As such, subject was a blocking variable to account for intersubject differences.
3. Results Scaling of the generic OpenSim model resulted in subjectspecific models with unique anthropometric parameters (Table 2). Trial durations for 1 lb hammer (0.81 ± 0.11 s, mean ± S.D) were significantly shorter than for 2 lb hammer (0.99 ± 0.22 s) (p < 0.01). Likewise, durations were shorter for the smaller hammer for cocking phase (0.34 ± 0.07 s and 0.44 ± 0.18s) (p < 0.05), for swing phase (0.07 ± 0.04 s and 0.12 ± 0.07 s) and strike/follow-through (0.39 ± 0.03 s and 0.43 ± 0.03 s) (both p < 0.01). Intersegmental moments varied throughout the nail strike motion both within and between phases for the 1 lb (Fig. 4) and 2 lb (Fig. 5) hammers. The larger hammer generally resulted in increased peak intersegmental moments at all upper extremity
Table 2 Scaled model segment masses. Mass [kg] (S.D.) Upper-extremity Clavicle Scapula Humerus Ulna Radius Hand
4.53 0.15 0.67 1.90 1.05 0.22 0.55
(0.89) (0.03) (0.13) (0.37) (0.21) (0.04) (0.11)
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Fig. 4. Intersegmental moments for (A) shoulder, (B) elbow and (C) wrist for 1 lb hammer. Grey denotes individual trials; black represents ±1 standard deviation from mean. Plane of elevation moment is positive counter-clockwise on right shoulder; other moments are positive for labeled direction. Moments are piece-wise linear length normalized to the average trial and phase durations for 1 lb hammers.
joints although the increase differed in terms of relative magnitude and significance of the effect (Fig. 6). Absolute measured moments and changes due to the heavier hammer were largest at the shoulder, followed by the elbow, and then the wrist. The largest moments were for shoulder plane of elevation and were somewhat smaller for shoulder elevation, while elbow flexion moments were substantially larger than other elbow moments for all phases of hammering. At the wrist during cocking and swinging phases, the largest moments were for ulnar deviation, while during strike/ follow-through, radial deviation tended to be relatively large with other moments also considerably increased. In general, increases in peak moments tended to be more pronounced in cocking and striking/follow-through phases as opposed to swinging phase. There was a trend toward larger relative increases in moments for elbow and wrist as compared to shoulder: 119% and 99% compared to 68% on average, respectively (Fig. 6). More specifically, at the shoulder the heavier hammer resulted in a significantly increased maximum elevation moment (by 9%) during the cocking (29.9 ± 2.5 vs. 27.5 ± 2.9 N-m, mean ± standard
error, p ¼ 0.0001) and strike/follow-through (89%) phases (55.8 ± 13.6 vs. 29.4 ± 2.8 N-m, p ¼ 0.002), while for internal rotation a significant increase (1.4 ± 0.1 vs. 0.9 ± 0.1 N-m, p ¼ 0.01) was found (53%) during the cocking phase. Likewise, the minimum moments for these joint angles (representing shoulder depression and external rotation, respectively) were also significantly different; the larger hammer resulted in increased shoulder depression moment during the strike phase (71.9 ± 10.6 vs. 49.7 ± 2.8 N-m, p ¼ 0.004) and increased external rotation moment (192%) in cocking phase (1.1 ± 0.1 vs. 0.4 ± 0.1 N-m, p < 0.0001). No significant differences were detected in peak shoulder elevation angle moments. At the elbow, the changes in peak moments were not as large in magnitude, but generally were more consistent across joint angles and demonstrated greater relative changes (Fig. 6B). With the heavier hammer, maximum elbow flexion moments across all phases were significantly greater (all p < 0.0001) by an average of 77% (e.g. for cocking, 14.5 ± 0.2 vs. 8.4 ± 0.1 N-m), while minimum elbow flexion moments were also significantly different with
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Fig. 5. Intersegmental moments for (A) shoulder, (B) elbow and (C) wrist for 2 lb hammer. Grey denotes individual trials; black represents ±1 standard deviation from mean. Plane of elevation moment is positive counter-clockwise on right shoulder; other moments are positive for labeled direction. Moments are piece-wise linear length normalized to the average trial and phase durations for 2 lb hammers.
mixed effect (p ¼ 0.03, p < 0.0001, p < 0.001, for phases respectively). For the cocking and swing phases, the peak minimum elbow flexion moments were in flexion direction, while for the strike/ follow-through phase the minimum moment was in extension. The peak maximum elbow pronation moment was significantly increased (average 120%) during cocking (0.5 ± 0.0 vs. 0.2 ± 0.0 Nm, p < 0.001) and strike/follow-through (2.4 ± 0.4 vs. 1.1 ± 0.1 N-m, p ¼ 0.03). Likewise, the peak minimum elbow pronation moment (i.e. supination) was significantly increased for all three phases by an average 143%) (1.7 ± 0.2 vs. 0.8 ± 0.1 N-m for strike/followthrough, p ¼ 0.001). The larger hammer caused a significant increase (138%) in elbow varus moment for the cocking phase (0.8 ± 0.1 vs. 0.3 ± 0.1 N-m, p ¼ 0.015) while the peak minimum varus (i.e. valgus) moment was significantly increased for all phases by an average 144% (e.g. for strike/follow-through, 7.4 ± 1.4 vs. 3.1 ± 0.3 N-m, p ¼ 0.04). When compared to the moments and changes at the shoulder and elbow, wrist peak moments and changes were generally small in magnitude, but for some components still demonstrated rather
large effect (Fig. 6C). The heavier hammer caused increased (by 99%) wrist flexion moment for strike/follow-through (1.9 ± 0.2 vs. 1.0 ± 0.1 N-m, p ¼ 0.0001), whereas the peak minimum moment (representing wrist extension) was significantly increased for all phases on average 83% (for strike/follow-through, 2.4 ± 0.2 vs. 1.3 ± 0.1 N-m, p < 0.001). Peak maximum ulnar deviation moment was significantly increased with the larger hammer during the swing (1.6 ± 0.2 vs. 0.6 ± 0.0 N-m, p < 0.001) and strike/followthrough phases (5.3 ± 0.5 vs. 2.7 ± 0.1 N-m, p < 0.0001). The peak minimum ulnar moments were in radial direction for all phases and were significantly increased (by average 108%) with 2 lb hammer (cocking phase: 2.5 ± 0.3 vs. 1.2 ± 0.1 N-m, and p < 0.001). The measured peak maximum wrist rotation moments were in the internal direction for the cocking and strike/follow-through phases and were significantly increased on average 226% with the larger hammer with large effect (for strike/follow-through, 2.3 ± 0.2 vs. 1.0 ± 0.1 N-m, p < 0.0001). Lastly, peak minimum wrist rotation moments were all negative (representing external rotation) for all phases and were significantly increased (74% on average) with the
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Fig. 6. Mean (±standard error) peak (A) shoulder, (B) elbow and (C) wrist intersegmental moments [N-m] for 1 and 2 pound hammers during phases of driving a nail. Significant differences denoted with * for p < 0.05 and ** for p < 0.01. Negative shoulder moments represent: elevation plane in counter-clockwise direction, depression, external rotation. Negative elbow moments represent: valgus, extension, supination. Negative wrist moments represent: radial deviation, extension, external rotation. (Note difference in scale for shoulder internal rotation and elbow varus and pronation moments).
large hammer (for strike/follow-through, 2.8 ± 0.3 vs. 1.4 ± 0.1 N-m, p ¼ 0.005). 4. Discussion This study investigated effects of increased hammer mass and moment of inertia on upper extremity joint kinetics during hammering of a nail. The motivation for this study is the hypothesis that larger hammers cause greater loading to the upper extremity musculoskeletal tissues. Results determined here that the larger hammer resulted in increased intersegmental joint moments for some components and some phases of the hammering motion support this hypothesis. Relatively large changes in intersegmental moments (50 to > 150%) were recorded when hammer moments of inertia were increased in the range of 113e150%. Additionally, trial time and phase durations were significantly greater (~10e30%) and greater variability (~50%) was demonstrated in these times for the larger hammer. Moments were generally smaller in absolute magnitude than reported in previous studies of upper extremity (Creveaux et al., 2013a,b; Rogowski et al., 2014; Slavens et al., 2010) but were similar to these studies and those of other joints, e.g. (Lerner et al., 2014), in that larger moments were calculated for primary degrees of freedom associated with generating the motion.
In addition, the hypothesis that a larger hammer would result in greater relative increases in moments for secondary degrees of freedom was supported. Taken together, the changes in moments and timing indicate greater loads on joints and musculoskeletal tissues for an extended period with the larger hammer. Subjects likely found the larger hammer more challenging to control as they required more time on average and demonstrated greater variability in duration for cocking and swinging phases than for the smaller hammer. Differences in duration might also have been related to novice subjects engaged in hammering. By definition, novices generate greater variance than more experienced subjects e.g. two subjects tended to generate larger moments in relation to others (Figs. 4 and 5). Subjects with less expertise likely experienced greater difficulty executing the coordinated motion across multiple joints required to generate large moments which resulted in the few differences found for shoulder moments, e.g. no differences were found at the shoulder during swing phase as subjects might have tended to let the hammer fall rather than actively extending the shoulder. In addition, the task was not representative of all hammering. This strategy and other potential differences described here could be ascertained in the future by measuring muscle activity during hammering. By assigning a task of striking a small nail and thus
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requiring novice subjects to optimize accuracy as well as force/ power, the resulting motion was of low acceleration when compared to a task where accuracy is not as paramount, such as striking a larger target, e.g. driving a stake of large cross-section where greater acceleration and force delivery to the target would be achieved. Such a task would likely result in greater loads imparted to the upper extremity for all phases, particularly when using a larger hammer, and might result in greater differences between the hammers. It is important to realize that for many of the secondary degrees of freedom where large relative changes in moments were found for some joints and components, the absolute changes in joint moments were small during cocking and swing phases, making the clinical significance of these results less likely and more difficult to assess. Even though the absolute increases were small, they were statistically significant for many components and there was a distinct trend towards increased moments when using the larger hammer that might be more pronounced for a different, more representative hammering task as mentioned above. Absolute increases were however larger for most joints and components during strike/follow through, indicating greater potential for injury during this phase. The relatively large increase for shoulder elevation during strike/follow-through indicates potential for higher concentric load placed on rotator cuff muscles and tendons, particularly supraspinatus, when hammer strikes a target. The increased depression moment during strike follow-through likely causes eccentric loading on rotator cuff, and to some extent deltoid, as these muscles experience lengthening contractions while slowing or arresting the hammer. Additionally, long head of biceps and its tendon would experience increased loading as it is partially responsible for aiding in humeral head depression while primarily responsible for the increased elbow flexion observed in all phases. Future studies to measure EMG of these and other muscles during hammering could clarify these potential injury mechanisms as previously performed for tennis racket swing (Giangarra et al., 1993). At the elbow during strike/follow through increased varusvalgus moments indicate greater tensile loads applied to radial and ulnar collateral ligaments, respectively, and also compressive loads to articular cartilage due to resulting force imbalance and inability of ligaments to fully resist the loading (Morrey and An, 1983). Increased elbow flexion moment results in greater loads applied to biceps tendon in all phases and especially with increased extension moment during strike/follow through when biceps could experience an eccentric contraction. At elbow, increased pronation and supination moments in strike/follow through require increased forces from pronators teres and quadratus and forearm supinator and biceps, respectively, thus increasing the potential for injuries to these muscles especially with repetitive use. For the wrist, increased radial and ulnar deviation moments during strike phase combined with greater flexion and extension moments would overload flexor carpi ulnaris and radialis as well as extensors carpi radialis longus, radialis brevis and ulnaris as commonly observed in clinical settings. The conjoined tendons for these muscles likely experience disorders in a manner similar to golf and tennis overuse injuries, i.e. radial and ulnar epicondylitis, due to eccentric contractions experienced when subjected to relatively large moments of changing direction during the short duration strike/follow through phase (Creveaux et al., 2013a,b; Knudson and Bahamonde, 1997). The trend of larger relative increases in moments for elbow and wrist when using greater mass hammers supports and agrees with the limited but similar findings of increased activity for forearm muscles as compared to upper arm and shoulder muscles when workers are required to use tools having more degrees of freedom which provide greater mechanical, physiological and psychological challenge (Fischer et al., 2009). However, in
comparing across joints, it is important to remember these increases were larger in a relative sense; absolute increases were larger for proximal joints. Because the hammers used in this study were of different mass, the overall magnitude of inertia differed accordingly, but remarkably the symmetries of inertia were nearly identical. The magnitude and symmetry of inertia for objects moved in space have been assumed to affect the intensity and patterning of muscle forces, respectively (Fitzpatrick et al., 2012). It would be interesting to investigate whether hammers with more varying magnitudes and particularly symmetries of inertia would require different motor control, and thus affect loads on upper extremity joints resulting in greater differences in kinematics and accelerations for both primary and especially secondary degrees of freedom. Additional work is necessary to understand hammering in these more representative, and likely higher loading situations. There are several limitations of this work related to the model and experimental task. First, although a generic model of the upper extremity was scaled to anthropometry of subjects, little information is available in literature to document the accuracy of this process. For lower-extremity, intersegmental forces and moments are usually normalized by body weight and product of body weight and height, respectively. Given the substantial inter-subject variability (Figs. 4 and 5) and the absence of similar scaling laws for upper-extremity, non-normalized moments were analysed statistically by blocking across subjects in order to account for intersubject variability. Next, although the interaction between the hammer and hand occurs as a distributed contact stress, by modeling all segments as rigid bodies as required for an inverse dynamics analysis, the force and moment between hammer and hand were assumed to be concentrated at a point on the hand and therefore calculated directly from hammer inertial properties and kinematics. In order to determine contact stresses between hammer and hand and at other joints, a deformable analysis, e.g. finite element analysis, would be required. Lastly, differences in handle size and shape were not considered here but have been shown to affect wrist and hand muscle forces and related load sharing (Rossi et al., 2012, Rossi et al., 2015). In conclusion, previous studies have suggested that mechanical loading is a risk factor for upper extremity injury (Buckle and Devereux, 2002). This study found increased intersegmental moments when using a larger hammer. Increases in joint moments when using a larger hammer were small in absolute magnitude for some joints, but there was a trend towards larger relative moments for the elbow and wrist, and particularly for secondary degrees of freedom such as elbow varus-valgus and pronation-supination and wrist rotation and radial-ulnar deviation. With the larger hammer, subjects required more time for the hammering motion and more time within phases with greater variability of the timing. These data and findings support previous understanding of the deleterious effects of hammer usage on upper extremity musculoskeletal tissues, and provide measurements of potential loads and describe mechanisms of the loadings. This information can potentially help explain an increased risk of developing cumulative trauma disorders due to repetitive hammer use and be used to inform future studies on hammering. Conflict of interest statement No author of this work has any financial or other conflict of interest. Acknowledgements This study was made possible and supported by equipment
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purchased under NSF-CBE-MRI 1337511. Sponsors had no role in: study design, data collection or analysis, interpretation of data, in manuscript writing or decision to submit the manuscript for publication. The authors thank Katherine Saul and Meghan Vidt for assistance with implementation of the MoBL-ARMs model. References Bernstein, N., 1967. The Coordination and Regulation of Movement. Pergamon Press, London. Buckle, P.W., Devereux, J.J., 2002. The nature of work-related neck and upper limb musculoskeletal disorders. Appl. Ergon. 33, 207e217. Chaffin, D.B., Andersson, G.B.J., 1999. Anthropometry in Occupational Biomechanics. In: Occupational Biomechanics. John Wiley and Sons, New York, NY. Cote, J.N., Feldman, A.G., Mathieu, P.A., Levin, M.F., 2008. Effects of fatigue on intermuscular coordination during repetitive hammering. Mot. Control 12, 79e92. Cote, J.N., Raymond, D., Mathieu, P.A., Feldman, A.G., Levin, M.F., 2005. Differences in multi-joint kinematic patterns of repetitive hammering in healthy, fatigued and shoulder-injured subjects. Clin. Biomech. 20, 581e590. , P., Che ze, L., Rogowski, I., 2013a. Joint Creveaux, T., Dumas, R., Hautier, C., Mace kinetics to assess the influence of the racket on a tennis player's shoulder. J. Sports Sci. Med. 12, 259e266. , P., Che ze, L., Rogowski, I., 2013b. Influence of racket Creveaux, T., Dumas, R., Mace polar moment on joint loads during tennis forehand drive. Comput. Methods Biomechancis Biomed. Eng. 16, 99e101. Delp, S.L., Anderson, F.C., Arnold, A.S., Loan, P., Habib, A., John, C.T., Guendelman, E., Thelen, D.G., 2007. OpenSim: open-source software to create and analyze dynamic simulations of movement. IEEE Trans. Biomed. Eng. 55, 1940e1950. Fischer, S.L., Wells, R.P., Dickerson, C.R., 2009. The effect of added degrees of freedom and handle type on upper limb muscle activity during simulated hand tool use. Ergonomics 52, 25e35. Fitzpatrick, P., Wagman, J.B., Schmidt, R.C., 2012. Alterations in movement dynamics in a tool-use task: the role of action-relevant inertial tool parameters. Z. fur Psychol. 220, 23e28. Giangarra, C.E., Conroy, B., Jobe, F.W., Pink, M., Perry, J., 1993. Electromyographic and cinematographic analysis of elbow function in tennis players using single- and double-handed backhand strokes. Am. J. Sports Med. 21, 394e399. Hammarskjold, E., Harms-Ringdahl, K., 1992. Effect of arm-shoulder fatigue on carpenters at work. Eur. J. Appl. Physiol. Occup. Physiol. 64, 402e409. Kahrs, B.A., Jung, W.P., Lockman, J.J., 2014. When does tool use become distinctively human? Hammering in young children. Child. Dev. 85, 1050e1061. Karkowski, W., Mattila, M., Quesada, P.M., Kwon, Y.G., James, D., 2002/2003. The effects of hammer design and hammering surface orientation on task performance. Occup. Ergon. 3, 45e54. Khan, A.A., Shakeb, M., 2009. Effect of posture, length of hammer-handle and frequency on discomfort. In: Developments in Agricultural and Industrial Ergonomics: General Studies, Vol. 1. Allied Publishers, Delhi. Knowlton, R.G., Gilbert, J.C., 1983. Ulnar deviation and short-term strength reductions as affected by a curve-handled ripping hammer and conventional claw hammer. Ergonomics 26, 171e179. Knudson, D., Bahamonde, R.E., 1997. Upper extremity angular kinematics of the one-handed backhand drive in tennis players with and without tennis elbow. Int. J. Sports Med. 18, 79e82.
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