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Three-dimensional lumbar segment kinetics of fast bowling in cricket
ARTICLE IN PRESS Journal of Biomechanics 42 (2009) 1616–1621
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Three-dimensional lumbar segment kinetics of fast bowling in cricket Rene´ E.D. Ferdinands a,, Uwe Kersting b,1, R.N. Marshall c,2 a b c
Discipline of Exercise, Health and Performance, The University of Sydney, East Street, PO Box 170, Lidcombe, NSW 1825, Australia SMI, Aalborg University, Fredrik Bajers Vej 7 E2-205, DK-9220, Aalborg, Denmark Eastern Institute of Technology, 501 Gloucester Street, Taradale, Hawke’s Bay 4142, New Zealand
a r t i c l e in f o
a b s t r a c t
Article history: Accepted 26 April 2009
Cricket fast bowlers have a high incidence of serious lumbar injuries, such as lesions in the pars interarticularis. Kinematic studies have shown that bowling actions with large shoulder counterrotation are associated with significantly higher incidences of lumbar injury. However, in bowling, there has been no calculation of the spinal loads, which are the causal mechanisms of such injuries. In this study, 21 fast bowlers (22.473.9 years) of premier grade level and above were tested using a threedimensional (3D) motion analysis system. The mean ball release speed was 31.972.8 m s 1 and ranged from 27.0 to 35.6 m s 1. Kinematics and kinetics were calculated for lumbar spine lateral bending, rotation, and flexion during the delivery and power phases of bowling. Power calculations were used to define the actuation of lumbar spine motion as either active or controlled. The actuation of the lumbar spine was complex, involving multiple sequences of active and controlled motion. In addition, lumbar spine loads were largest during the power phase when the ground reaction forces were highest. In conclusion, the dynamic loads and the cyclical nature of their application when the spine is positioned near its end range of motion may be significant factors of injury to this region. In addition, the lumbar spine in bowling has to vigorously flex, laterally bend and rotate simultaneously in a complex interdependent sequence of actuation patterns. Therefore, any technical change to reduce injury susceptibility needs to consider the mechanics of whole body coordination and timing. & 2009 Elsevier Ltd. All rights reserved.
Keywords: Cricket Bowling Back injury Biomechanics Kinematics Kinetics
1. Introduction Fast bowling in cricket is a vigorous activity, which requires the lumbar spine to flex, laterally bend and rotate in a short period of time to produce ball speeds up to 45 m s 1 (Burnett et al., 1998), while absorbing large vertical ground reaction forces (3.8–6.4 times body weight) at front foot impact (Bartlett et al., 1996). In addition, as the segment at the nexus between the upper and lower body, the lumbar spine has to withstand the loading generated by the contractions of its own muscle actuators, while absorbing or transferring the reactive loads from all the other segments that ‘flow’ to it by means of the kinetic linked chain mechanism (Winter and Robertson, 1978; Putnam, 1993). In general, lumbar vertebra or disc injury is the result of accumulated trauma from the cyclical application of a nontraumatic spinal load or the application of a sustained spinal load so that there is a slow degradation of tissue failure tolerance in the region (McGill, 1997). Experiments on the lumbar vertebrae of Corresponding author. Tel.: +61 2 9351 9776; fax: +61 2 9351 9204.
E-mail addresses: [email protected] (R.E.D. Ferdinands), [email protected] (U. Kersting), [email protected] (R.N. Marshall). 1 Tel.: +45 9635 8094; fax: +45 9815 4008. 2 Tel.: +64 6 974 8000x5422; fax: +64 6 974 8910. 0021-9290/$ - see front matter & 2009 Elsevier Ltd. All rights reserved. doi:10.1016/j.jbiomech.2009.04.035
porcine models, which have similar tissue failure tolerance thresholds to that of human tissue, show that compressive and shear spinal loading cause stresses and strains on the vertebral structures such as the pars interarticularis, which may eventually lead to fracture (Yingling and McGill, 1999). Repetitive mechanical loading of the thoracolumbar spine is considered a major causal factor associated with pars interarticularis stress fractures of cricket fast bowlers (Elliott, 2000; Engstrom and Walker, 2007), which can progress to spondylolisthesis (Burnett et al., 1995; Elliott et al., 1993). Bowlers may also sustain various abnormalities of the intervertebral discs such as reduced height, degeneration, bulging and herniation as a result of bowling activity (Elliott, 2000). Shoulder counter-rotation, which is the maximum clockwise rotation of a shoulder alignment vector projected on the horizontal plane, has been associated with such back injuries (Elliott, 2000; Portus et al., 2004), but no causal mechanism has been identified. Therefore, from the perspective of understanding spinal injury mechanisms, it is essential to calculate the lumbar spine loads, which has not been previously done for bowling. The purpose of this study is to investigate the magnitude and temporal characteristics of three-dimensional (3D) lumbar spine kinetics in fast bowling. To relate these loads to the configuration of lumbar spine during the phases of the bowling action, the kinematics were also measured. This will provide information on
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the dynamic loading patterns on the lumbar spine in bowling, which has implications for bowling technique and the identification of potential injury mechanisms.
2. Methods 2.1. Sample and experimental protocol Twenty-one fast bowlers (22.473.9 years) from the New Zealand premier grade club level and above were recruited. All the subjects provided informed consent and reported that they were free from any injury or physical dysfunction, which may have affected bowling performance at the time of data collection. Ethical approval for this study was obtained from the research ethics committees of the University of Auckland and the University of Waikato.
2.2. Data collection The trials were performed in a biomechanics laboratory, which permitted a 20 m length run-up. A 8-camera EVa Motion Analysis System (Motion Analysis Corporation Ltd., USA) was used to capture three-dimensional motion (240 Hz) and force plate (960 Hz) data on six trials for each bowler, while front and rear foot contact was made on two Bertec 6090 force plates. Each subject was instructed to bowl at maximum effort and pitch the ball within a ‘good length’ area demarcated by two white lines 15 and 19 m from the stumps at the bowler’s end. Subjects had also to rate their performance from 0 to 10 using an analogue performance scale. The capture volume encompassed the back foot contact, front foot contact, ball release and follow-through phases of the bowling action. The EVa system was calibrated according to the manufacturer’s recommendations resulting in a residual error of marker position of less than 2 mm.
2.3. Kinematic marker model Motion analysis capture was performed on each subject wearing a full body marker set comprising forty-five 15 mm spherical markers, which were attached to bony landmarks (Ferdinands, 2004). Markers were located on the left and right sides of the body except for markers placed half-way between the posterior superior iliac spines (mid-PSIS), and on the 7th cervical vertebrae, supra-sternal notch, T10, sternum and the head. Most marker positions were chosen to define joint centres to delimit segment lengths as listed in de Leva (1996). Exceptions were the position of the shoulder, mid-trunk, hip markers, and cricket ball. The positions of the anterior superior iliac spine (ASIS), mid-PSIS, and greater trochanter markers were as recommended by Bell et al. (1990), who used the positions of these markers to calculate the hip joint centres. All other joint centres were calculated as the average position between two markers placed either medially and laterally or anteriorly and posteriorly on the joint. Local segment coordinate systems of a fifteen-segment rigid body model were derived using the standard approach of Grood and Suntay (1983), and embedded at the proximal ends of their respective segments. The lumbar spine segment (LSS) was defined as a single segment having its proximal end located half-way between the hip joint centres at the level of L5/S1. The distal endpoint was located at the mid-point between the markers on the supra-sternal notch and T10. The hip segment was defined using a local pelvic coordinate system comprising the sacrum and both ASIS markers. A local coordinate system was also defined at the distal end of the LSS at the T10 level. The motion of the model in three-dimensional space was described using the xyz-coordinates of the origin of the local coordinate systems and the zyx-Euler angle sequence of local coordinate systems with respect to the global laboratory coordinate system. A full set of three-dimensional linear and angular kinematic data were calculated. A recursive fourth-order low-pass Butterworth filter was used to smooth the kinematic data (Winter, 2005). The cut-off frequencies were determined from the Fast Fourier Transform analysis functions in Mathematica (Version 5.2).
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2.5. Data analysis The fastest ball that was also given a subjective player rating of 7 or above was selected for analysis for each bowler. Net joint torque was calculated about the proximal end of the LSS for lateral bending, flexion/extension and rotation. Right lateral bending, flexion and left rotation were defined as positive. Lumbar motion was characterised as either active or controlled (Winter and Robertson, 1978; Ferdinands, 2004). Active motion was defined when the joint torque and segmental angular velocity were in the same sense and corresponded to power generation or net concentric muscle action. Controlled motion was defined when the joint torque and segmental angular velocity were in the opposite sense and corresponded to power absorption or net eccentric muscle action. LSS kinematic and kinetic data for all subjects were pooled and ensemble averaged (mean 795% CI) over the delivery stride and power periods. The delivery stride phase was defined from back foot contact to front foot contact. The power phase was defined from front foot contact to ball release. These phases were timenormalised individually and expressed as percentage 795% confidence interval (CI). In addition, relative LSS flexion-extension, rotation and lateral angular displacement data were expressed as a percentage of the ranges of motion about their respective principal axes. The ranges of motion were estimated by adding the range of motion of each spinal level from S1 to T10 based on the data from McGill (2007). Therefore, the LSS flexion-extension, rotation and bending ranges of motion were estimated as 691, 191 and 531, respectively. Note that the calculation of percentage range of LSS rotation was based on the relative orientation of the joint coordinate systems at the distal end with respect to the proximal end of the LSS. Pelvis rotation was based on the projected rotation of the pelvic coordinate system on the laboratory coordinate system with respect to the X-axis, which was pointing along the line of pitch to the batter’s end.
3. Results The mean ball release speed was 31.972.8 m s from 27.0 to 35.6 m s 1.
1
and ranged
3.1. Flexion–extension From an examination of the power–torque curves in Fig. 1, there were four distinct periods of LSS movement in the flexion/ extension plane during the delivery stride phase (Table 1). From a slightly extended position of the LSS at back foot contact ( 5.572.01), there was a short period of small controlled flexion (0–7%) before a slightly longer period of active flexion (8–28%: 3.472.01). This was followed by a controlled extension phase from 28% to 83%, when 10.777.81 of extension was controlled with a mean flexion torque of 160.6774.3 N m. Towards the end of the phase, there was a small active extension, which led into the power phase.
2.4. Kinetic model The three-dimensional kinematic data were imported into a fifteen-segment inverse dynamics model of the cricket bowler (Ferdinands, 2004). The model was designed in the Mathematica Mechanical Systems Pack (Version 5.2, Wolfram Research Ltd.), which analyses the kinematics and kinetics of spatial rigid body mechanisms by solving the equations of motion using the Newtonian–Lagrange multiplier iterative method (Haug, 1989; Dynamic Modelling, 1995).
Fig. 1. Ensemble average 70.95 CI of LSS flexion torque and power during the delivery stride phase (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; flexion torque defined as positive).
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Table 1 Lumbar spine segment movement during the delivery stride phase.
Phase (%) 0 10 20 30 40 50 60 70 80 90 100
FlexionExtension Flexion
Actuation Type Controlled
Flexion
Active
Extension
Controlled
Rotation
Actuation Type
Counterrotation
Controlled
Active Left Rotation
Extension
Active
Lateral Bending
Actuation Type
Left Lateral Bending
Controlled
Controlled
(82–100% period). A negative percentage indicates that the T10 level vertebra was rotated more to the left relative to the lumbosacral joint. There was a mean maximum torque of 266.77140.0 Nm at 22%. 3.3. Lateral bending
Fig. 2. Ensemble average 70.95 CI of LSS flexion torque and power during the power phase. The power was positive throughout the phase and peaked at 34%, occurring temporally close to the time of maximum horizontal and vertical ground reaction forces (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; flexion torque defined as positive).
Fig. 2 shows that during the subsequent power phase, there was a large flexion torque (412.17174.2 Nm), which actuated an active LSS flexion of 38.278.01 (55.4711.6% range of motion) (Table 2) with a mean angular velocity of 310.3793.81/s. There was a mean maximum flexion torque of 742.07442.4 Nm at 34%.
An evaluation of Fig. 5 shows that during delivery stride, the LSS was in controlled left lateral bending (Table 1). From a small initial LSS right lateral bending angle (10.672.81) at back foot contact, there was a mean controlled lateral bending of 23.774.21 to the left (44.777.9% range of motion) until front foot contact with a mean angular velocity of 121.9747.81/s. Fig. 6 shows that LSS was actively laterally bending during most of the power phase (Table 2). The LSS did actively laterally bend a further 15.7711.31 to the left with a mean angular velocity of 148.2796.61/s. At ball release, the LSS had used up 74.3714.6% of the available range of lateral bending motion. The mean maximum left lateral bending torque was 533.87 122.1 Nm at 24%. 3.4. Ground reaction forces Maximum horizontal and vertical ground reaction forces were 18587309.4 N and 31267346.9 N at 25% and 27% of the power phase, respectively. These maxima corresponded approximately with the maximum joint muscle powers for flexion (34%), rotation (22%), and lateral bending (23%).
3.2. Rotation From an evaluation of Fig. 3, there were three distinct periods of LSS rotation during the delivery phase (Table 1). From an initial pelvis rotation of 56.1712.71 to the left, there was a small controlled pelvis counter-rotation (3.572.41) during approximately 0–36%. This was followed by a pelvis rotation of 31.2712.61 to the left (mean angular velocity 60.1758.21/s) that was active from 37% to 76% and then controlled until front foot contact. There was a large increase in the percentage use of the available range of LSS rotation to the left during the period of controlled left rotation (77–100%) from 49.279.7% to 107.978.8%. During the subsequent power phase, the pelvis rotated a further 22.9710.31 to the left (mean angular velocity of 74.5774.31/s). Fig. 4 shows that the LSS was in controlled left rotation during this entire phase (Table 2). The LSS was using up approximately the entire range of available rotation (102.4712.8%) during the entire period when the power was greater than 100 W (0–50%). However, the LSS used smaller percentages of available left rotation during the lower power periods: 36.5720.6% (51–81% period) and 20.6712.6%
4. Discussion The aim of this study was to investigate and identify the magnitude and temporal characteristics of three-dimensional lumbar spine kinetics in fast bowling. Kinematics was used to describe the loading positions of the lumbar spine during bowling. As the spine rotates through its range of motion, the vertebrae, discs and ligaments are placed under more strain, which increases the risk of tissue injury (White and Panjabi, 1990; Panjabi, 1992; Ranson et al., 1998), particularly if there is large cyclical loading (Solomonow et al., 2000). Therefore, lumbar spine motion was also quantified as the percentage use of available range of motion. These calculations were used to investigate the mechanics of the lumbar spine in bowling and identify potential injury mechanisms. 4.1. LSS extension–flexion In the modern game, most bowlers are classified as front-on (Portus et al., 2004; Ferdinands, 2004). Hence, any LSS extension
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Table 2 Lumbar spine segment movement during the power phase.
Phase (%)
FlexionExtension
Actuation Type
Rotation
Actuation Type
0 10 20 30 40 50 60 70 80 90 100
Flexion
Active
Left Rotation
Controlled
Lateral Bending L Lateral Bending
Actuation Type Controlled
Left Lateral Bending
Active
Right Lateral Bending
Active
Fig. 3. Ensemble average 70.95 CI of LSS rotation torque and power during the delivery phase (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; left rotation torque defined as positive).
Fig. 5. Ensemble average 70.95 CI of LSS lateral bending torque and power during the delivery phase. Maximum power was at 23% occurring temporally close to the time of maximum horizontal and vertical ground reaction forces (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; right lateral bending torque defined as positive).
Fig. 4. Ensemble average 70.95 CI of LSS rotation and torque power during the power phase. The power was negative, indicating that the motion was controlled. Maximum power was at 22% occurring temporally close to the time of maximum horizontal and vertical ground reaction forces (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; left rotation torque defined as positive).
Fig. 6. Ensemble average 70.95 CI of LSS lateral bending torque and power during the power period (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; right lateral bending torque defined as positive).
during the delivery phase is intended to increase the range of flexion during the subsequent power phase, which may lead to increased ball release speed. Although the LSS was actually
slightly flexed at back foot contact and was subject to a series of four actuation types (Table 1), there was relatively small lumbar spine extension or flexion during the delivery phase. There was
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only 10.777.81 of LSS extension, which was mostly controlled by a net joint muscle flexion torque from 30% to 80% of the delivery phase. This is consistent with the theory that there are interdependent mechanical interactions in a linked segment system (Putnam, 1993) and motion of the LSS can be caused by the movement coordination patterns of other body segments. During the power phase, there was actively actuated LSS flexion (Fig. 2). Although the bowlers only used a moderate percentage (55.4711.6%) of the available range of LSS flexion, the mean flexion angular velocity of the LSS motions was high, indicating there may be considerable more use of the available range of LSS flexion in the follow-through phase due to momentum. There were also large flexion torques during the power phase. Therefore, the lumbar spine had to withstand a large load at high angular velocity generated by the trunk flexor musculature, and potentially operate towards the end of its range of available flexion. Lumbar vertebrae under these conditions are placed under large stresses, and are prone to fracture particularly under cyclical loading (McGill, 1997; Solomonow et al., 2000). Although such conditions are regarded as the causal mechanisms of tissue injury, they have not been investigated in any assessment of injury risk in bowling. 4.2. Lateral bending of LSS During all of the delivery phase and the first 10% of the power phase, the LSS lateral bending to the left was controlled (Tables 1 and 2) by the lateral bending torque to the right acting on the LSS during this period (Figs. 5 and 6). However, during the remainder of the power phase, there was a powerful active left lateral bending of the LSS (Table 2), using 74.3714.6% of the available range of motion. This means that both left lateral bending and flexion of the LSS were actively actuated by the lower lumbar spinal musculature during the power phase. This corresponds to kinematics studies, which have found that trunk flexion angular velocity is associated with ball release speed (Bartlett et al., 1996). However, several researchers have found that lateral bending velocity of the trunk is a risk factor for low-back disorders (Marras and Granata, 1997). High active lateral bending to the left may therefore promote lumbar injury in fast bowlers, particularly as a high percentage of the available range of motion is used. This may also explain why the incidence of lumbar stress injuries occurs predominantly on the opposite side of the bowling arm (Elliott et al., 1992; Ranson et al., 2005).
ultimate compressive strength of vertebral tissue (Aultman et al., 2004). As the fast bowlers in this study had large flexion torques, which load the spine in compression, techniques that produce large rotation torques during the power phase may further increase the risk of lumbar injury. Although bowlers had a high rotation angular velocity of the LSS to the left during the power phase, this was produced in the opposite sense to the LSS torque, which acted to the right. However, if all the body segments in bowling are part of an interacting dynamic system, then the effects of the action–reaction forces and torques on each segment are propagated throughout the kinetics chain and ‘experienced’ as mechanical inputs to every other segment in the chain (Putnam, 1993). In this way, the motion of each segment influences every other segment in the chain. As in the case of LSS rotation to the left in the power phase, it is therefore entirely possible for a segment to move without any muscle actuation, but as the result of the sum of interacting reactive moment inputs from the other segments in the chain. Similarly, LSS rotation to the right (hip counterrotation) was entirely controlled by a net joint rotation torque to the left.
4.4. Ground reaction forces The largest horizontal and vertical ground reaction forces occurred between 20% and 40% of the power phase. The iterative nature of the power calculations means that the magnitude of the ground reaction forces are carried through the kinetic chains from the most distal to the most proximal leg segments where they are summed at the base of the lumbar spine (Skotte, 2001; Winter, 2005). Hence, the LSS loads for flexion–extension, lateral bending and rotation were also at their highest during this time (Fig. 2). In addition, large ground reaction forces occur when LSS angular velocities are high in all lumbar planes of motion. Therefore, the large dynamic loads on the lumbar spine are due to (i) the high ground reaction forces, (ii) the powerful muscle action that actuates or controls the motion of this relatively heavy segment, and (iii) the high angular velocities of the LSS. With high LSS torques and angular velocities, the dynamic loading on the spine during the power phase was correspondingly high, producing maximum LSS power values in all planes of motion, and placing the lumbar spine at increased risk of injury. In addition, this risk may be further increased if the repeated absorption of ground reaction forces by the lumbar spine occurs when it is near its anatomical end range of motion.
4.3. LSS rotation After an initial controlled counter-rotation period during the delivery phase, there was a large range of LSS rotation to the left with high angular velocity, which is comparable to LSS rotation in baseball pitching and javelin throwing (Pappa et al., 1985; Mero et al., 1994). However, the power data suggests that there were two actuation phases of LSS rotation (Tables 1 and 2). There was an active period of relatively low power (30–70% of the delivery stride phase), which was followed by a lengthy period of controlled LSS rotation to the left until the end of the power phase. During this controlled period, the LSS was using approximately its full range of available rotation, and there were a high torque and power during the early part of the power phase when ground reaction forces were high (Fig. 4). From an injury perspective, rotation or axial torques may reduce the effective compressive strength of the spine. This is particularly the case when the LSS has increased stiffness when operating towards the end of its range of motion (Panjabi, 1992). Based on tests of porcine cervical spines, additional torsion compromises the
4.5. Limitations There are several limitations with this study. The anthropometric values for the lumbar spine are based on regression equations from de Leva (1996). These are approximations because the lumbar spine is not a rigid object, and the accuracy of the predictive equations would depend on the properties such as the size and fat content of the abdominal region. This leads to errors in joint torque calculations, which multiply upwards through the kinetic chain, because inverse dynamics is sensitive to differences in segment inertial values. The motion of the LSS was measured using a skin-based marker system, and not bone pins, which are more accurate markers for segment end-points. The calculation of percentage of available LSS range of motion was based on population averages and do not consider the typically large individual variations in such data. In addition, it was assumed that the fast bowlers had a ‘‘normal’’ range of spine motion to validate the estimation of the relative lumbar range of motion.
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5. Conclusion The purpose of this study was to investigate the LSS kinetics in bowling. It was found that the LSS has to vigorously flex, laterally bend and rotate simultaneously in a complex interdependent sequence of actuation patterns. This implies that any intended change in the kinematics of the lumbar spine to reduce the susceptibility to injury in this region should not be made without considering the coordinated rhythm of the entire segmental sequence. The LSS was also subject to very high loading, particularly during the power phase. As bowlers typically bowl thousands of balls each season, and vertebral tissue failure tolerance threshold decreases with cyclical loading (McGill, 1997), the dynamics loads and the cyclical nature of their application may be significant factors of injury to this region (Yingling and McGill, 1999), particularly when the LSS is configured close to the end of its range of motion. Future research should investigate the relationship between injury and spinal loading magnitudes and patterns, and extend this analysis to the follow-through phase.
Conflict of interest statement There are no conflicts of interest to disclose for the authors.
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Three-dimensional lumbar segment kinetics of fast bowling in cricket Rene´ E.D. Ferdinands a,, Uwe Kersting b,1, R.N. Marshall c,2 a b c
Discipline of Exercise, Health and Performance, The University of Sydney, East Street, PO Box 170, Lidcombe, NSW 1825, Australia SMI, Aalborg University, Fredrik Bajers Vej 7 E2-205, DK-9220, Aalborg, Denmark Eastern Institute of Technology, 501 Gloucester Street, Taradale, Hawke’s Bay 4142, New Zealand
a r t i c l e in f o
a b s t r a c t
Article history: Accepted 26 April 2009
Cricket fast bowlers have a high incidence of serious lumbar injuries, such as lesions in the pars interarticularis. Kinematic studies have shown that bowling actions with large shoulder counterrotation are associated with significantly higher incidences of lumbar injury. However, in bowling, there has been no calculation of the spinal loads, which are the causal mechanisms of such injuries. In this study, 21 fast bowlers (22.473.9 years) of premier grade level and above were tested using a threedimensional (3D) motion analysis system. The mean ball release speed was 31.972.8 m s 1 and ranged from 27.0 to 35.6 m s 1. Kinematics and kinetics were calculated for lumbar spine lateral bending, rotation, and flexion during the delivery and power phases of bowling. Power calculations were used to define the actuation of lumbar spine motion as either active or controlled. The actuation of the lumbar spine was complex, involving multiple sequences of active and controlled motion. In addition, lumbar spine loads were largest during the power phase when the ground reaction forces were highest. In conclusion, the dynamic loads and the cyclical nature of their application when the spine is positioned near its end range of motion may be significant factors of injury to this region. In addition, the lumbar spine in bowling has to vigorously flex, laterally bend and rotate simultaneously in a complex interdependent sequence of actuation patterns. Therefore, any technical change to reduce injury susceptibility needs to consider the mechanics of whole body coordination and timing. & 2009 Elsevier Ltd. All rights reserved.
Keywords: Cricket Bowling Back injury Biomechanics Kinematics Kinetics
1. Introduction Fast bowling in cricket is a vigorous activity, which requires the lumbar spine to flex, laterally bend and rotate in a short period of time to produce ball speeds up to 45 m s 1 (Burnett et al., 1998), while absorbing large vertical ground reaction forces (3.8–6.4 times body weight) at front foot impact (Bartlett et al., 1996). In addition, as the segment at the nexus between the upper and lower body, the lumbar spine has to withstand the loading generated by the contractions of its own muscle actuators, while absorbing or transferring the reactive loads from all the other segments that ‘flow’ to it by means of the kinetic linked chain mechanism (Winter and Robertson, 1978; Putnam, 1993). In general, lumbar vertebra or disc injury is the result of accumulated trauma from the cyclical application of a nontraumatic spinal load or the application of a sustained spinal load so that there is a slow degradation of tissue failure tolerance in the region (McGill, 1997). Experiments on the lumbar vertebrae of Corresponding author. Tel.: +61 2 9351 9776; fax: +61 2 9351 9204.
E-mail addresses: [email protected] (R.E.D. Ferdinands), [email protected] (U. Kersting), [email protected] (R.N. Marshall). 1 Tel.: +45 9635 8094; fax: +45 9815 4008. 2 Tel.: +64 6 974 8000x5422; fax: +64 6 974 8910. 0021-9290/$ - see front matter & 2009 Elsevier Ltd. All rights reserved. doi:10.1016/j.jbiomech.2009.04.035
porcine models, which have similar tissue failure tolerance thresholds to that of human tissue, show that compressive and shear spinal loading cause stresses and strains on the vertebral structures such as the pars interarticularis, which may eventually lead to fracture (Yingling and McGill, 1999). Repetitive mechanical loading of the thoracolumbar spine is considered a major causal factor associated with pars interarticularis stress fractures of cricket fast bowlers (Elliott, 2000; Engstrom and Walker, 2007), which can progress to spondylolisthesis (Burnett et al., 1995; Elliott et al., 1993). Bowlers may also sustain various abnormalities of the intervertebral discs such as reduced height, degeneration, bulging and herniation as a result of bowling activity (Elliott, 2000). Shoulder counter-rotation, which is the maximum clockwise rotation of a shoulder alignment vector projected on the horizontal plane, has been associated with such back injuries (Elliott, 2000; Portus et al., 2004), but no causal mechanism has been identified. Therefore, from the perspective of understanding spinal injury mechanisms, it is essential to calculate the lumbar spine loads, which has not been previously done for bowling. The purpose of this study is to investigate the magnitude and temporal characteristics of three-dimensional (3D) lumbar spine kinetics in fast bowling. To relate these loads to the configuration of lumbar spine during the phases of the bowling action, the kinematics were also measured. This will provide information on
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the dynamic loading patterns on the lumbar spine in bowling, which has implications for bowling technique and the identification of potential injury mechanisms.
2. Methods 2.1. Sample and experimental protocol Twenty-one fast bowlers (22.473.9 years) from the New Zealand premier grade club level and above were recruited. All the subjects provided informed consent and reported that they were free from any injury or physical dysfunction, which may have affected bowling performance at the time of data collection. Ethical approval for this study was obtained from the research ethics committees of the University of Auckland and the University of Waikato.
2.2. Data collection The trials were performed in a biomechanics laboratory, which permitted a 20 m length run-up. A 8-camera EVa Motion Analysis System (Motion Analysis Corporation Ltd., USA) was used to capture three-dimensional motion (240 Hz) and force plate (960 Hz) data on six trials for each bowler, while front and rear foot contact was made on two Bertec 6090 force plates. Each subject was instructed to bowl at maximum effort and pitch the ball within a ‘good length’ area demarcated by two white lines 15 and 19 m from the stumps at the bowler’s end. Subjects had also to rate their performance from 0 to 10 using an analogue performance scale. The capture volume encompassed the back foot contact, front foot contact, ball release and follow-through phases of the bowling action. The EVa system was calibrated according to the manufacturer’s recommendations resulting in a residual error of marker position of less than 2 mm.
2.3. Kinematic marker model Motion analysis capture was performed on each subject wearing a full body marker set comprising forty-five 15 mm spherical markers, which were attached to bony landmarks (Ferdinands, 2004). Markers were located on the left and right sides of the body except for markers placed half-way between the posterior superior iliac spines (mid-PSIS), and on the 7th cervical vertebrae, supra-sternal notch, T10, sternum and the head. Most marker positions were chosen to define joint centres to delimit segment lengths as listed in de Leva (1996). Exceptions were the position of the shoulder, mid-trunk, hip markers, and cricket ball. The positions of the anterior superior iliac spine (ASIS), mid-PSIS, and greater trochanter markers were as recommended by Bell et al. (1990), who used the positions of these markers to calculate the hip joint centres. All other joint centres were calculated as the average position between two markers placed either medially and laterally or anteriorly and posteriorly on the joint. Local segment coordinate systems of a fifteen-segment rigid body model were derived using the standard approach of Grood and Suntay (1983), and embedded at the proximal ends of their respective segments. The lumbar spine segment (LSS) was defined as a single segment having its proximal end located half-way between the hip joint centres at the level of L5/S1. The distal endpoint was located at the mid-point between the markers on the supra-sternal notch and T10. The hip segment was defined using a local pelvic coordinate system comprising the sacrum and both ASIS markers. A local coordinate system was also defined at the distal end of the LSS at the T10 level. The motion of the model in three-dimensional space was described using the xyz-coordinates of the origin of the local coordinate systems and the zyx-Euler angle sequence of local coordinate systems with respect to the global laboratory coordinate system. A full set of three-dimensional linear and angular kinematic data were calculated. A recursive fourth-order low-pass Butterworth filter was used to smooth the kinematic data (Winter, 2005). The cut-off frequencies were determined from the Fast Fourier Transform analysis functions in Mathematica (Version 5.2).
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2.5. Data analysis The fastest ball that was also given a subjective player rating of 7 or above was selected for analysis for each bowler. Net joint torque was calculated about the proximal end of the LSS for lateral bending, flexion/extension and rotation. Right lateral bending, flexion and left rotation were defined as positive. Lumbar motion was characterised as either active or controlled (Winter and Robertson, 1978; Ferdinands, 2004). Active motion was defined when the joint torque and segmental angular velocity were in the same sense and corresponded to power generation or net concentric muscle action. Controlled motion was defined when the joint torque and segmental angular velocity were in the opposite sense and corresponded to power absorption or net eccentric muscle action. LSS kinematic and kinetic data for all subjects were pooled and ensemble averaged (mean 795% CI) over the delivery stride and power periods. The delivery stride phase was defined from back foot contact to front foot contact. The power phase was defined from front foot contact to ball release. These phases were timenormalised individually and expressed as percentage 795% confidence interval (CI). In addition, relative LSS flexion-extension, rotation and lateral angular displacement data were expressed as a percentage of the ranges of motion about their respective principal axes. The ranges of motion were estimated by adding the range of motion of each spinal level from S1 to T10 based on the data from McGill (2007). Therefore, the LSS flexion-extension, rotation and bending ranges of motion were estimated as 691, 191 and 531, respectively. Note that the calculation of percentage range of LSS rotation was based on the relative orientation of the joint coordinate systems at the distal end with respect to the proximal end of the LSS. Pelvis rotation was based on the projected rotation of the pelvic coordinate system on the laboratory coordinate system with respect to the X-axis, which was pointing along the line of pitch to the batter’s end.
3. Results The mean ball release speed was 31.972.8 m s from 27.0 to 35.6 m s 1.
1
and ranged
3.1. Flexion–extension From an examination of the power–torque curves in Fig. 1, there were four distinct periods of LSS movement in the flexion/ extension plane during the delivery stride phase (Table 1). From a slightly extended position of the LSS at back foot contact ( 5.572.01), there was a short period of small controlled flexion (0–7%) before a slightly longer period of active flexion (8–28%: 3.472.01). This was followed by a controlled extension phase from 28% to 83%, when 10.777.81 of extension was controlled with a mean flexion torque of 160.6774.3 N m. Towards the end of the phase, there was a small active extension, which led into the power phase.
2.4. Kinetic model The three-dimensional kinematic data were imported into a fifteen-segment inverse dynamics model of the cricket bowler (Ferdinands, 2004). The model was designed in the Mathematica Mechanical Systems Pack (Version 5.2, Wolfram Research Ltd.), which analyses the kinematics and kinetics of spatial rigid body mechanisms by solving the equations of motion using the Newtonian–Lagrange multiplier iterative method (Haug, 1989; Dynamic Modelling, 1995).
Fig. 1. Ensemble average 70.95 CI of LSS flexion torque and power during the delivery stride phase (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; flexion torque defined as positive).
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Table 1 Lumbar spine segment movement during the delivery stride phase.
Phase (%) 0 10 20 30 40 50 60 70 80 90 100
FlexionExtension Flexion
Actuation Type Controlled
Flexion
Active
Extension
Controlled
Rotation
Actuation Type
Counterrotation
Controlled
Active Left Rotation
Extension
Active
Lateral Bending
Actuation Type
Left Lateral Bending
Controlled
Controlled
(82–100% period). A negative percentage indicates that the T10 level vertebra was rotated more to the left relative to the lumbosacral joint. There was a mean maximum torque of 266.77140.0 Nm at 22%. 3.3. Lateral bending
Fig. 2. Ensemble average 70.95 CI of LSS flexion torque and power during the power phase. The power was positive throughout the phase and peaked at 34%, occurring temporally close to the time of maximum horizontal and vertical ground reaction forces (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; flexion torque defined as positive).
Fig. 2 shows that during the subsequent power phase, there was a large flexion torque (412.17174.2 Nm), which actuated an active LSS flexion of 38.278.01 (55.4711.6% range of motion) (Table 2) with a mean angular velocity of 310.3793.81/s. There was a mean maximum flexion torque of 742.07442.4 Nm at 34%.
An evaluation of Fig. 5 shows that during delivery stride, the LSS was in controlled left lateral bending (Table 1). From a small initial LSS right lateral bending angle (10.672.81) at back foot contact, there was a mean controlled lateral bending of 23.774.21 to the left (44.777.9% range of motion) until front foot contact with a mean angular velocity of 121.9747.81/s. Fig. 6 shows that LSS was actively laterally bending during most of the power phase (Table 2). The LSS did actively laterally bend a further 15.7711.31 to the left with a mean angular velocity of 148.2796.61/s. At ball release, the LSS had used up 74.3714.6% of the available range of lateral bending motion. The mean maximum left lateral bending torque was 533.87 122.1 Nm at 24%. 3.4. Ground reaction forces Maximum horizontal and vertical ground reaction forces were 18587309.4 N and 31267346.9 N at 25% and 27% of the power phase, respectively. These maxima corresponded approximately with the maximum joint muscle powers for flexion (34%), rotation (22%), and lateral bending (23%).
3.2. Rotation From an evaluation of Fig. 3, there were three distinct periods of LSS rotation during the delivery phase (Table 1). From an initial pelvis rotation of 56.1712.71 to the left, there was a small controlled pelvis counter-rotation (3.572.41) during approximately 0–36%. This was followed by a pelvis rotation of 31.2712.61 to the left (mean angular velocity 60.1758.21/s) that was active from 37% to 76% and then controlled until front foot contact. There was a large increase in the percentage use of the available range of LSS rotation to the left during the period of controlled left rotation (77–100%) from 49.279.7% to 107.978.8%. During the subsequent power phase, the pelvis rotated a further 22.9710.31 to the left (mean angular velocity of 74.5774.31/s). Fig. 4 shows that the LSS was in controlled left rotation during this entire phase (Table 2). The LSS was using up approximately the entire range of available rotation (102.4712.8%) during the entire period when the power was greater than 100 W (0–50%). However, the LSS used smaller percentages of available left rotation during the lower power periods: 36.5720.6% (51–81% period) and 20.6712.6%
4. Discussion The aim of this study was to investigate and identify the magnitude and temporal characteristics of three-dimensional lumbar spine kinetics in fast bowling. Kinematics was used to describe the loading positions of the lumbar spine during bowling. As the spine rotates through its range of motion, the vertebrae, discs and ligaments are placed under more strain, which increases the risk of tissue injury (White and Panjabi, 1990; Panjabi, 1992; Ranson et al., 1998), particularly if there is large cyclical loading (Solomonow et al., 2000). Therefore, lumbar spine motion was also quantified as the percentage use of available range of motion. These calculations were used to investigate the mechanics of the lumbar spine in bowling and identify potential injury mechanisms. 4.1. LSS extension–flexion In the modern game, most bowlers are classified as front-on (Portus et al., 2004; Ferdinands, 2004). Hence, any LSS extension
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Table 2 Lumbar spine segment movement during the power phase.
Phase (%)
FlexionExtension
Actuation Type
Rotation
Actuation Type
0 10 20 30 40 50 60 70 80 90 100
Flexion
Active
Left Rotation
Controlled
Lateral Bending L Lateral Bending
Actuation Type Controlled
Left Lateral Bending
Active
Right Lateral Bending
Active
Fig. 3. Ensemble average 70.95 CI of LSS rotation torque and power during the delivery phase (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; left rotation torque defined as positive).
Fig. 5. Ensemble average 70.95 CI of LSS lateral bending torque and power during the delivery phase. Maximum power was at 23% occurring temporally close to the time of maximum horizontal and vertical ground reaction forces (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; right lateral bending torque defined as positive).
Fig. 4. Ensemble average 70.95 CI of LSS rotation and torque power during the power phase. The power was negative, indicating that the motion was controlled. Maximum power was at 22% occurring temporally close to the time of maximum horizontal and vertical ground reaction forces (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; left rotation torque defined as positive).
Fig. 6. Ensemble average 70.95 CI of LSS lateral bending torque and power during the power period (Power, solid line; torque, dashed line; zero reference lines, dashpot lines; right lateral bending torque defined as positive).
during the delivery phase is intended to increase the range of flexion during the subsequent power phase, which may lead to increased ball release speed. Although the LSS was actually
slightly flexed at back foot contact and was subject to a series of four actuation types (Table 1), there was relatively small lumbar spine extension or flexion during the delivery phase. There was
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only 10.777.81 of LSS extension, which was mostly controlled by a net joint muscle flexion torque from 30% to 80% of the delivery phase. This is consistent with the theory that there are interdependent mechanical interactions in a linked segment system (Putnam, 1993) and motion of the LSS can be caused by the movement coordination patterns of other body segments. During the power phase, there was actively actuated LSS flexion (Fig. 2). Although the bowlers only used a moderate percentage (55.4711.6%) of the available range of LSS flexion, the mean flexion angular velocity of the LSS motions was high, indicating there may be considerable more use of the available range of LSS flexion in the follow-through phase due to momentum. There were also large flexion torques during the power phase. Therefore, the lumbar spine had to withstand a large load at high angular velocity generated by the trunk flexor musculature, and potentially operate towards the end of its range of available flexion. Lumbar vertebrae under these conditions are placed under large stresses, and are prone to fracture particularly under cyclical loading (McGill, 1997; Solomonow et al., 2000). Although such conditions are regarded as the causal mechanisms of tissue injury, they have not been investigated in any assessment of injury risk in bowling. 4.2. Lateral bending of LSS During all of the delivery phase and the first 10% of the power phase, the LSS lateral bending to the left was controlled (Tables 1 and 2) by the lateral bending torque to the right acting on the LSS during this period (Figs. 5 and 6). However, during the remainder of the power phase, there was a powerful active left lateral bending of the LSS (Table 2), using 74.3714.6% of the available range of motion. This means that both left lateral bending and flexion of the LSS were actively actuated by the lower lumbar spinal musculature during the power phase. This corresponds to kinematics studies, which have found that trunk flexion angular velocity is associated with ball release speed (Bartlett et al., 1996). However, several researchers have found that lateral bending velocity of the trunk is a risk factor for low-back disorders (Marras and Granata, 1997). High active lateral bending to the left may therefore promote lumbar injury in fast bowlers, particularly as a high percentage of the available range of motion is used. This may also explain why the incidence of lumbar stress injuries occurs predominantly on the opposite side of the bowling arm (Elliott et al., 1992; Ranson et al., 2005).
ultimate compressive strength of vertebral tissue (Aultman et al., 2004). As the fast bowlers in this study had large flexion torques, which load the spine in compression, techniques that produce large rotation torques during the power phase may further increase the risk of lumbar injury. Although bowlers had a high rotation angular velocity of the LSS to the left during the power phase, this was produced in the opposite sense to the LSS torque, which acted to the right. However, if all the body segments in bowling are part of an interacting dynamic system, then the effects of the action–reaction forces and torques on each segment are propagated throughout the kinetics chain and ‘experienced’ as mechanical inputs to every other segment in the chain (Putnam, 1993). In this way, the motion of each segment influences every other segment in the chain. As in the case of LSS rotation to the left in the power phase, it is therefore entirely possible for a segment to move without any muscle actuation, but as the result of the sum of interacting reactive moment inputs from the other segments in the chain. Similarly, LSS rotation to the right (hip counterrotation) was entirely controlled by a net joint rotation torque to the left.
4.4. Ground reaction forces The largest horizontal and vertical ground reaction forces occurred between 20% and 40% of the power phase. The iterative nature of the power calculations means that the magnitude of the ground reaction forces are carried through the kinetic chains from the most distal to the most proximal leg segments where they are summed at the base of the lumbar spine (Skotte, 2001; Winter, 2005). Hence, the LSS loads for flexion–extension, lateral bending and rotation were also at their highest during this time (Fig. 2). In addition, large ground reaction forces occur when LSS angular velocities are high in all lumbar planes of motion. Therefore, the large dynamic loads on the lumbar spine are due to (i) the high ground reaction forces, (ii) the powerful muscle action that actuates or controls the motion of this relatively heavy segment, and (iii) the high angular velocities of the LSS. With high LSS torques and angular velocities, the dynamic loading on the spine during the power phase was correspondingly high, producing maximum LSS power values in all planes of motion, and placing the lumbar spine at increased risk of injury. In addition, this risk may be further increased if the repeated absorption of ground reaction forces by the lumbar spine occurs when it is near its anatomical end range of motion.
4.3. LSS rotation After an initial controlled counter-rotation period during the delivery phase, there was a large range of LSS rotation to the left with high angular velocity, which is comparable to LSS rotation in baseball pitching and javelin throwing (Pappa et al., 1985; Mero et al., 1994). However, the power data suggests that there were two actuation phases of LSS rotation (Tables 1 and 2). There was an active period of relatively low power (30–70% of the delivery stride phase), which was followed by a lengthy period of controlled LSS rotation to the left until the end of the power phase. During this controlled period, the LSS was using approximately its full range of available rotation, and there were a high torque and power during the early part of the power phase when ground reaction forces were high (Fig. 4). From an injury perspective, rotation or axial torques may reduce the effective compressive strength of the spine. This is particularly the case when the LSS has increased stiffness when operating towards the end of its range of motion (Panjabi, 1992). Based on tests of porcine cervical spines, additional torsion compromises the
4.5. Limitations There are several limitations with this study. The anthropometric values for the lumbar spine are based on regression equations from de Leva (1996). These are approximations because the lumbar spine is not a rigid object, and the accuracy of the predictive equations would depend on the properties such as the size and fat content of the abdominal region. This leads to errors in joint torque calculations, which multiply upwards through the kinetic chain, because inverse dynamics is sensitive to differences in segment inertial values. The motion of the LSS was measured using a skin-based marker system, and not bone pins, which are more accurate markers for segment end-points. The calculation of percentage of available LSS range of motion was based on population averages and do not consider the typically large individual variations in such data. In addition, it was assumed that the fast bowlers had a ‘‘normal’’ range of spine motion to validate the estimation of the relative lumbar range of motion.
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5. Conclusion The purpose of this study was to investigate the LSS kinetics in bowling. It was found that the LSS has to vigorously flex, laterally bend and rotate simultaneously in a complex interdependent sequence of actuation patterns. This implies that any intended change in the kinematics of the lumbar spine to reduce the susceptibility to injury in this region should not be made without considering the coordinated rhythm of the entire segmental sequence. The LSS was also subject to very high loading, particularly during the power phase. As bowlers typically bowl thousands of balls each season, and vertebral tissue failure tolerance threshold decreases with cyclical loading (McGill, 1997), the dynamics loads and the cyclical nature of their application may be significant factors of injury to this region (Yingling and McGill, 1999), particularly when the LSS is configured close to the end of its range of motion. Future research should investigate the relationship between injury and spinal loading magnitudes and patterns, and extend this analysis to the follow-through phase.
Conflict of interest statement There are no conflicts of interest to disclose for the authors.
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