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Western Washington University
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WWU Masters Thesis Collection
Spring 2015
Abdominal Fatigue and Lower Extremity
Kinematics During a Drop Landing in Females.
Eryn N. Murphy
Western Washington University, [email protected]
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Murphy, Eryn N., "Abdominal Fatigue and Lower Extremity Kinematics During a Drop Landing in Females." (2015). WWU Masters
Thesis Collection. Paper 405.
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Abdominal Fatigue and Lower Extremity Kinematics During a Drop Landing in Females.
By
Eryn Murphy
Accepted in Partial Completion
of the Requirements for the Degree
Masters of Science
Kathleen L. Kitto, Dean of Graduate School
ADVISORY COMMITTEE
Chair, Dr. Jun G. San Juan
Dr. Lorrie R. Brilla
Dr. David N. Suprak
MASTER’S THESIS
In presenting this thesis in partial fulfillment of the requirements for a master’s degree at
Western Washington University, I grant to Western Washington University the non-exclusive
royalty-free right to archive, reproduce, distribute, and display the thesis in any and all forms,
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allowed without my written permission.
Eryn Murphy
May 6th, 2015
Abdominal Fatigue and Lower Extremity Kinematics During a Drop Landing in Females.
A Thesis
Presented to
The Faculty of
Western Washington University
In Partial Fulfillment
Of the Requirements for the Degree
Master of Science
By
Eryn Murphy
May 2015
Abstract
This purpose of this study was to investigate the acute effects of abdominal fatigue on
lower-extremity risk factors of anterior cruciate ligament (ACL) injury. Kinematic variables
associated with ACL injury include, decreased knee flexion, increased knee valgus, increased hip
flexion and increased hip internal rotation. Additionally, increased quadriceps activation relative
to hamstring activation is a risk factor, since the quadriceps place strain on the anterior knee.
Subjects were instructed to sit on the ground, knees bent to 90°, feet secured and their
trunk supported by a wedge at 40°. The 40° posture was sustained for as long and as accurately
as possible. Pre- and post-test included 3 consecutive drop landings from 30.5 cm. Kinematic
data was analyzed from the second pre-test landing, as well as the first and the second post-test
landing. Muscle activation data was analyzed from the second pre-test landing, and the second
post-test landing.
Statistical analysis was carried out using a one-way repeated measures analysis of
variance (ANOVA) for kinematic variables. After fatigue, there was a significant effect of time
on knee flexion (p<0.05). Post-hoc analysis revealed a significant increase between pre-test and
both post-test values (p=0.012 and p=0.006, post-test 1 and post-test 2 respectively). There was a
non-significant effect of time on knee valgus, hip flexion and hip internal rotation. A paired
sample t-test was performed to analyze significant differences in muscle activation patterns after
abdominal fatigue. There was a significant increase in quadriceps activation (p<0.05), and an
associated non-significant increase in quadriceps:hamstring ratio. In conclusion, abdominal
fatigue may increase the risk of ACL injury in college aged females by altering biomechanical
variables associated with ACL injury.
iv
Acknowledgements
I would like to thank the faculty of Western Washington University’s Physical Education,
Health and Recreation department for helping raise the rigor of my work to prepare me for the
completion of this thesis. I would like to extend a special thank you to Dr. Jun San Juan, Dr.
Dave Suprak and Dr. Lorrie Brilla, as members of my thesis committee, for holding me to the
highest standards. Over the last two years, each one of you have taken extensive time out of your
hectic schedules to listen, mentor, and advise me through this process, and for that I am forever
grateful.
To Dr. Jun San Juan, I thank you for your commitment to the quality of my work and the
numerous hours you spent to help craft this idea into a reality. Our conversations forced me to
look at this field in a new light, and pushed my work to higher heights. Foremost, I thank you for
your never-ending patience. I know without a shadow of a doubt that all of my future work will
be remarkably more professional because of your guidance.
I extend a special thank you to Erik Hummer, Tylre Arens and Joseph Cordell for
assisting me during the data collection process. Without you, I would have no thesis to submit.
Not only did you all help maintain the integrity of my study, but you provided me with
friendship and sanity through the entire process.
Finally, I owe an incredible thank you to my friends and family for supporting me
through this process. More than being patient with me through this adventure, you provided me
with unwavering support, compassion and encouragement from the beginning. I am forever
indebted to you, without you, none of this would have been possible.
v
Table of Contents
Abstract ……..…………………………………...………………………………… iv
Acknowledgements…....…………………………………………………………… v
List of figures and tables…….……………………………………………………... ix
Chapter 1; The Problem and Its Scope
Introduction….……………………………………………………………………... 1
Purpose of the Study……………………………………………………………..… 3
Experimental Hypothesis…………………………………………………………... 3
Significance of the Study…………………………………………………………... 3
Limitations of the Study………………………………………………………….…4
Definition of Terms………………………………………………………................5
Chapter II; Review of Literature
Introduction………………………………………………………………………… 7
Review of Literature……………………………………………………………..… 8
Anatomy of the trunk………………………………………………………. 8
Anterior musculature…………………………………………….… 9
Fatigue………………………………………………………………………10
Biomechanics of fatigue…………………………………………………… 10
Trunk fatigue……………………………………………………………….. 11
General trunk fatigue and lower extremities……………………… 11
General influence of fatigue on the lower extremities……………... 13
Kinematic risk factors for ACL injury……………………………………... 14
Lower extremity kinematics……………………………………………….. 15
vi
Lower extremity muscle activation patterns……………………………….. 16
Fatiguing protocols….……………………………………………………... 17
Drop landing protocols…………………………………………………….. 19
Summary…………………………………………………………………… 20
Chapter III; Methodology
Introduction… ………………………………………………………………………22
Description of study population……………………………………………………. 22
Design of study…………………………………………………………………….. 23
Data collection procedures…………………………………………………………. 23
Instrumentation…………………………………………………………….. 23
Surface electromyography………………..………………………... 23
Three dimensional motion analysis ……....………...……………... 24
Measurement techniques and testing procedures…………………………... 26
Data processing ………………………………………………………….… 32
Data analysis……………………………………………………………….. 33
Chapter IV; Results and Discussion
Introduction………………………………………………………………………… 34
Subject characteristics……………………………………………………………… 34
Results……………………………………………………………………………… 34
Joint kinematics……………………………………………………………. 35
Muscle activation…………………………………………………………... 38
Discussion….………………………………………………………………………. 42
Summary…………………………………………………………………………… 50
vii
Chapter V; Summary, Conclusions, and Recommendations………………………………. 51
Summary…………………………………………………………………………… 51
Conclusions.………………………………………………………………………... 51
Recommendations………………………………………………………………….. 52
Future research……………………………………………………………………... 52
Practical applications………………………………………………………. 52
References………………………………………………………………………….. 54
Appendices…………………………………………………………………………. 66
viii
Lists of Figures and Tables
Figure 1.
Noraxon dual electrodes…….…………………………………………..
24
Figure 2a.
Lateral view of subject instrumentation ………………………………..
25
Figure 2b.
Posterior view of subject instrumentation ………………………..…….
25
Figure 2c.
Anterior view of subject instrumentation ………………………………
26
Figure 3.
Laboratory data collection lay out ……………………………….……..
30
Figure 4.
Abdominal fatigue subject set up ……………………………………….
31
Figure 5.
Abdominal fatigue hold posture ………………………………………...
32
Figure 6.
Maximum knee flexion graph …………………………………………...
36
Figure 7.
Maximum knee valgus graph ……………………………………………
36
Figure 8.
Maximum hip flexion graph …………………………………………….
37
Figure 9.
Maximum hip internal rotation graph …………………………………… 38
Figure 10.
Mean muscle activation graph …………………………………………... 39
Figure 11.
Peak muscle activation graph ……………………………………………. 39
Figure 12.
Peak quadriceps activation graph ………………………………………… 40
Figure 13.
Mean quadriceps to hamstring comparison graph ……………..………... 40
Figure 14.
Peak quadriceps to hamstring comparison graph ………………………... 41
Figure 15.
Peak quadriceps to hamstring ratio graph …...…………..………………
Figure 16.
Mean quadriceps to hamstring ratio graph ……………………………...… 42
Table 1.
Electrode placement and reference contractions …………………………. 27
ix
41
Chapter I
The Problem and Its Scope
Introduction
Research trends demonstrate that abdominal strength decreases the risk of ACL injury
(Leetun, Ireland, Willson, Ballantyne, & Davis, 2004). However, much of this research occurs in
highly controlled situations, monitoring trunk displacement after perturbation while the lower
extremities are immobilized (Zazulak, Hewett, Reeves, Goldberg, & Cholewicki, 2007), or while
incorporating elements of axial rotation (Ng, Parnianpour, Richardson, & Kippers, 2003).
Similarly, current research induces lower extremity fatigue (Kernozek, Torry, & Iwasaki, 2008;
Thomas, McLean & Palmieri-Smith, 2010), or aerobic fatigue (Quammen et al., 2012)
measuring a dependent variable at the lower extremities. There is limited research showing the
acute implications of fatigued abdominals on a distal and dynamic, multi-joint event such as a
drop landing (Walsh, Boling, McGrath, Blackburn, & Padua, 2012).
Females have a significantly increased risk for non-contact anterior cruciate ligament
(ACL) injury compared to their male counterparts (Arendt, Agel, & Dick, 1999; Arendt & Dick,
1995). Reportedly, females are 4- to 6-times more likely to sustain an ACL injury than males
(Arendt & Dick, 1995; Myer, Ford, Foss, Liu, Nick & Hewett, 2009). The anatomical differences
between genders contribute greatly to the injury and movement pattern discrepancies observed in
sport and activity (Kernozek et al., 2008). For example, females tend to have wider hips,
increasing the angle relative to their patella and tibia (Myer, Ford, & Hewett, 2005; Zazulak et
al., 2005), a combination increasing the degree of knee valgus. These anatomical differences are
well substantiated, and indicative of increased risk of ACL injury (Blackburn & Padua, 2008).
Movements including deceleration, land and change in direction are identified as
common movements leading to ACL injury (Boden, Dean, Feagin & Garret, 2000). Many ACL
injuries are non-contact in origin. As such, there are many biomechanical factors which increase
the risk of ACL injury, including knee valgus (Boden et al., 2009; Hewett et al., 2005), decreased
knee flexion (Walsh et al., 2012) increased hip flexion (Boden et al., 2009; Kernozek et al.,
2008) and increased hip internal rotation (McLean, Huang & Bogert, 2005).
Additionally, females demonstrate a different muscle recruitment pattern during dynamic
tasks compared to men, which may be indicative of knee injury. Myer et al. (2009) compared the
quadriceps to hamstring ratio (Q:H) of ACL injured female athletes, healthy age-matched
females and male control subjects during concentric knee extension and flexion strength at 300
degrees per second. Compared to male control subjects, injured females had a decreased knee
flexor strength of 15%, while female controls were not different from male controls in knee
flexor strength. However, injured females did not differ in quadriceps strength compared to male
controls, and interestingly, female controls demonstrated decreased quadriceps strength
compared to males controls (Myer, Ford, Foss, Liu, Nick & Hewett, 2009). These differences
identify an increase in Q:H to be an indicator for ACL injury (Cowling & Steele, 2000).
However, it is not clear the impact that abdominal fatigue has on risk factors of ACL
injury in women. Considering that it appears women are predisposed to ACL injury, based not
only on anatomical differences, but also in muscle recruitment patterns, it is important to identify
additional risk factors so that risk of injury can be minimized. With such great anatomical
connection between the abdominal musculature and the lower extremities (Blackburn & Padua,
2008) through the pelvic girdle (Foch, 2014; Norris, 1993; Willson, Dougherty, Ireland, &
2
Davis, 2005), it is expected that abdominal fatigue will alter the kinematics observed at the hip
and knee joints.
Purpose of the Study
The purpose of this study is to demonstrate the relationship between the anterior trunk
and the lower extremities, specifically helping to identify additional risk factors for anterior
cruciate ligament injury.
Experimental Hypothesis
The hypothesis is that abdominal fatigue will result in an increased degree of knee
valgus, decreased degree of knee flexion, increased degree of hip flexion and increased degree of
hip internal rotation during drop landing in females. Additionally, abdominal fatigue will result
in an increased quadriceps to hamstring activation ratio, as well as decreased gluteus maximus
activation, activation patterns consistent with the above kinematic variables.
Significance of the Study
Females show significantly higher anterior cruciate ligament (ACL) injury rates than
their male counterparts (Arendt et al., 1999; Arendt & Dick, 1995; Myer et al.; 2009). This is
likely due to the increased knee valgus forces and coupled hip abduction forces observed in
females during landing asks (Boden et al., 2009; Hewett et al., 2005). These biomechanical
differences between genders are based in anatomical differences; females tend to have wider
hips, narrower shoulders and a lower center of gravity compared to males (Smith et al., 2012;
Zazulak et al., 2005) as well differing muscular activation patterns between the knee extensors
3
and flexors (Myer et al., 2009). Zazulak et al. (2007) demonstrated that abdominal strength
predicted the risk of ACL injury in college athletes. While there is an understanding that
abdominal strength or lack thereof is related to ACL injury (Leetun et al., 2004; Olson, 2009;
Zazulak et al., 2007), it has not been determined how acute abdominal fatigue effects kinematic
risk factors of ACL injury.
Limitations of the Study

Due to the nature of the fatiguing protocol, subjects may recruit the hip flexors as
well as the posterior musculature of the trunk in addition to the trunk flexors,
more directly affecting the hip kinematics during the drop landing than isolated
trunk fatigue.

During the fatiguing protocol, subjects were encouraged to maintain the 40°
posture “as accurately and for as long as possible”. However, the posture that was
maintained through the fatigue protocol was not objectively controlled for.

The females recruited for this study had no history of ACL injury, and as such,
results can only be applied to a healthy college-aged females.

Fitness level and daily activities were not controlled for, consequently females of
varying athletic history likely showed different landing patterns.
4
Definition of Terms
Abduction
Movement of a limb away from the midline of the body. This movement
occurs in the frontal plane.
Adduction
Movement of a limb toward the midline of the body. This movement
occurs in the frontal plane.
Central Fatigue
A resulting insufficiency at the muscle due to failures at or above the
spine in the central nervous system. Typically occurs prior to peripheral
fatigue (Allen et al., 2008; Macintosh & Rassier, 2002; Vøllestad,
1997).
Distal
Located further away from the center of the body (i.e. the knee is distal
to the hip).
Extension
The increase in the angle of a joint. Extension at the knee and hip occurs
in the sagittal plane (Whiting & Zernicke, 2008).
Fatigue
As measured in this study: the inability to sustain a 40° isometric trunk
flexion task.
Flexion
The decrease in the angle of a joint. Flexion at the knee and hip occurs
in the sagittal plane (Whiting & Zernicke, 2008).
Frontal Plane
The frontal plane is a vertical plane which separates the body into
“front” and “back” aspects.
Inferior
To be located below a landmark (i.e. the ankle is inferior to the knee).
Kinematics
Measurements of the joints concerned with the motion of the body,
without reference to the forces that cause the movement (Whiting &
Zernicke, 2008)
5
Kinetics
Assessment of movement with regard to the forces acting on the body
(Whiting & Zernicke, 2008).
Kinetic Chain
Series and links of muscles, fascia, bones and ligaments working in
synchrony to transfer, produce, attenuate or store energy.
Peripheral Fatigue Resulting insufficiency at the muscle due to imbalances of key minerals
or substrates, leading to inadequate contractive abilities within the
muscle (Allen et al., 2008; MacIntosh et al., 2012; Macintosh & Rassier,
2002; Sahlin et al., 1998; Westerblad, Allen, & Lannergren, 2002;
Westerblad et al., 1998; Westerblad, Lee, Lannergren, & Allen, 1991).
Proximal
Located nearer the center of the body (i.e. the knee is proximal to the
ankle).
Sagittal Plane
The sagittal plane is a vertical plane which separates the body into left
and right aspects.
Superior
To be located above a landmark (i.e. the head is superior to the
abdomen).
Transverse Plane
The transverse plane is a horizontal plane which separates the body into
upper and lower aspects.
Valgus
In relation to this study, valgus refers to the displacement of the knee
joint toward the midline of the body.
Varus
In relation to this study, varus refers to the displacement of the knee
joint away from the midline of the body.
6
Chapter II
Review of Literature
Introduction
The human body is built of kinetic chains; series and links of muscles, fascia, bones and
ligaments working in synchrony to transfer, produce, attenuate or store energy. Recently,
research has shifted to investigate how proximal segments in these chains influence the abilities
or the performance of distal segments within the same kinetic chain (Hart, Kerrigan, Fritz, &
Ingersoll, 2009; Helbostad et al., 2010; Kulas, Hortobágyi, & DeVita, 2010; Olson, 2009;
Quammen et al., 2012; Sparto, Parnianpour, Reinsel, & Simon, 1997; Zazulak et al., 2007).
While localized fatigue is well understood as the inability of a muscle to perform a given
contraction (Allen, Lamb, & Westerblad, 2008; Dingwell, Joubert, Diefenthaeler, & Trinity,
2008; Kellis & Kouvelioti, 2009; MacIntosh, Holash, & Renaud, 2012; Macintosh & Rassier,
2002; Ng et al., 2003; Sahlin, Tonkonogi, & Soderlund, 1998; Vøllestad, 1997; Walsh et al.,
2012; Westerblad, Allen, Bruton, Andrade, & Lännergren, 1998), the impact of fatigue on a
distal segment within the same kinetic chain is less investigated.
Fatigue can fall into two categories: central and peripheral fatigue (Allen et al., 2008;
Macintosh & Rassier, 2002; Vøllestad, 1997). Central fatigue references the shortcomings
occurring at the central nervous system, and the consequent limitations at the muscle. Peripheral
fatigue is a direct comment to the environment of the muscle, peripheral nerves or the
neuromuscular junction (Macintosh & Rassier, 2002). Often, peripheral fatigue is due to
imbalances of key minerals or substrates, leading to inadequate contractive abilities within the
muscle (Allen et al., 2008; MacIntosh et al., 2012; Macintosh & Rassier, 2002; Sahlin et al.,
1998; Westerblad, Allen, & Lannergren, 2002; Westerblad et al., 1998; Westerblad, Lee,
7
Lannergren, & Allen, 1991). Under the majority of circumstances, central fatigue occurs prior to
the onset of peripheral fatigue (Allen et al., 2008; Macintosh & Rassier, 2002; Vøllestad, 1997),
as such, most research focuses on the effects of peripheral fatigue (Corin, Strutton, & McGregor,
2005; Coventry, O’Connor, Hart, Earl, & Ebersole, 2006; Dedering, Roos af Hjelmsäter, Elfving,
Harms-Ringdahl, & Németh, 2000; Dingwell et al., 2008; Hart et al., 2009; Helbostad et al.,
2010; Kellis & Kouvelioti, 2009; Kernozek et al., 2008; Ng et al., 2003; Olson, 2009; Page &
Descarreaux, 2012; Quammen et al., 2012; A. L. Rodacki, Fowler, & Bennett, 2001; Sparto et
al., 1997).
Fatigue of varying sources can alter the movement patterns of the lower extremities both
by muscle activation differences (Bellew & Fenter, 2006; Davidson et al., 2009; Helbostad et al.,
2010; Kavanagh, Morrison, & Barrett, 2006; Lin et al., 2009; Petrella, Kim, Tuggle, Hall, &
Bamman, 2005; Pline, Madigan, Nussbaum, & Grange, 2005) and joint kinematic relationships
(De Luca, 1984; Ng et al., 2003; Rodacki et al., 2001; Sparto et al., 1997). Similarly, fatigue can
reduce achievements in balance tasks related to core stabilization (Hart et al., 2009; Helbostad et
al., 2010; Kulas et al., 2010). With this understanding, this study aims to demonstrate the
relationship of abdominal fatigue with lower-extremity kinematics and muscle activation patterns
of the lower extremity.
Review of Literature.
Anatomy of the trunk. The trunk connects the upper and lower extremities, as well as
supports the head and neck. Specifically, the trunk contains the skeletal portions of the body
from the external occipital protuberances of the skull, down to the iliac crest and sacrum,
including the rib cage. Therefore, it includes also the muscles of the abdomen and back (Brown,
8
Ward, Cook, & Lieber, 2011; S. M. McGill, 1996). Muscles of the trunk play a vital role in
controlling posture and stability in both dynamic and static tasks (Park, Song, Park, Park, Bae, &
Lim, 2014).
Anterior musculature. Of primary concern, the muscles of the abdomen include the
internal and external obliques, the transverse abdominus, and rectus abdominus. As a whole, the
abdominal muscles work together to flex the spine in the frontal and sagittal plane. They also
provide stability to the spine by helping to increase intra-abdominal pressure (Brown et al., 2011;
S. M. McGill, 1996).
The internal oblique lies deep to the external oblique, and superficial to the transverse
abdominus. The fibers of the internal and external obliques run perpendicular to each other. The
internal oblique originates at the lumbar fascia, the lateral portion of the inguinal ligament and
anterior portion of the iliac crest. The insertion point of the internal oblique includes the costal
margin, aponeurosis of the rectus sheath and the pubic crest. The primary role of the internal
oblique is to aid in forced expiration and raising intra-abdominal pressure. The external obliques
originate at the 5th through 12th ribs and run to the anterior iliac crest and abdominal aponeurosis.
They play a large role in lateral flexion of the trunk as well as rotation and flexion (Brown et al.,
2011; S. M. McGill, 1996).
The transverse abdominis is deep to the internal oblique. Its fibers run transversely
around the abdomen, originating at the lateral third of the inguinal ligament, anterior portion of
the iliac crest, the last six ribs and inserts at the centrally located linea alba, xiphoid process, and
pubic crest. The transverse abdominis largely helps to provide stability to both the pelvis and the
ribs (Richardson et al., 2002). In comparison, the rectus abdominis primarily flexes the lumbar
spine. It originates at the crest of the pubis, and inserts at ribs 5-7, and the xiphoid process. It is
9
separated medially by the linea alba, and transversely by the tendinous intersections, which
separate the rectus abdominis into 8 separate portions (Brown et al., 2011; S. M. McGill, 1996).
Fatigue. Musculoskeletal fatigue is the inability to perform a given contraction despite
sufficient absolute strength (Allen et al., 2008; Dingwell et al., 2008; Kellis & Kouvelioti, 2009;
MacIntosh et al., 2012; Macintosh & Rassier, 2002; Ng et al., 2003; Sahlin et al., 1998;
Vøllestad, 1997; Walsh et al., 2012; Westerblad et al., 1998). Fatigue may be of central or
peripheral in origin. Central fatigue refers to restrictions at or above the spinal cord, where the
central nervous system is the limiting factor. Occurring secondary to central fatigue, peripheral
fatigue reflects an insufficiency within the muscle, the peripheral nerves or the neuromuscular
junction (Macintosh & Rassier, 2002). Underlying mechanisms of peripheral fatigue echo the
inability to generate ample Adenosine Triphosphate (Allen et al., 2008). Hypotheses of
peripheral fatigue include increased concentration of cellular inorganic phosphate (Allen et al.,
2008; Cady, Jones, Lynn, & Newham, 1989; Takagi, Y., Shuman, H., & Goldman, Y. E., 2004;
Westerblad et al., 2002), decreased sensitivity or availability of calcium (Fitts, 1994; Westerblad
et al., 1991), and rate of myosin light chain phosphorylation (Allen et al., 2008; Bottinelli &
Reggiani, 2000; MacIntosh et al., 2012; Sahlin et al., 1998).
Biomechanics of fatigue. Physiological mechanisms of fatigue lead to biomechanical
alterations (De Luca, 1984; Ng et al., 2003). For example, Ng, et al., concluded that coupling
torques were slightly decreased during fatiguing axial rotation in both sagittal and coronal
planes. Decreases in torque production (Rodacki et al., 2001; Sparto et al., 1997) occur as a
result of muscular fatigue. Similarly, fatigue leads to alterations of myoelectric activity; median
power frequency decreases in fatiguing muscles as measured by surface electromyography
(Dedering et al., 2000; Dingwell et al., 2008; Ng et al., 2003; Olson, 2009; Page & Descarreaux,
10
2012; Shultz, Nguyen, Leonard, & Schmitz, 2009). Depending on the fatiguing protocol, joint
moments and kinematics are also altered by muscular fatigue (Coventry, O' Connor, Hart, Earl &
Ebersole, 2006; Horita, Komi, Nicol, & Kyröläinen, 1999; Kernozek, Torry, Van Hoof, Cowley,
& Tanner, 2008).
In a study investigating the kinematic differences during a drop landing task between
men and women caused by muscular fatigue, Kernozek et al. (2008) discovered that significant
kinematic alterations occurred in women, indicative of ACL injury mechanisms. After
performing parallel squats until fatigue, women showed an increase in hip flexion and knee
valgus. Effects of fatigue reflected in kinematic differences were also demonstrated by Coventry
et al. (2006). Ten male subjects performed a series of lower extremity tasks including maximal
countermovement jumps, and squats until exhaustion. Results demonstrated an increase in hip
and knee flexion at initial contact during single leg landings (Coventry et al., 2006).
Trunk fatigue. Abdominal fatigue can be induced by either isokinetic or isometric
contractions (Corin et al., 2005). Isometric fatigue protocols deliver the most valid assessment of
fatigue in comparison to isokinetic protocols of testing and inducing fatigue (Corin et al., 2005).
Most research confirms the influence of aerobic or localized fatigue on lower extremities
(Dingwell et al., 2008; Helbostad et al., 2010; Kellis & Kouvelioti, 2009; Kernozek et al., 2008;
Kulas et al., 2010; Quammen et al., 2012; Rodacki, Fowler, & Bennett, 2002; Rodacki et al.,
2001; Sparto et al., 1997). Although recent research has confirmed correlations between the
trunk and lower extremities during activities (Hart et al., 2009; Kulas et al., 2010), specific
correlations between abdominal fatigue and lower extremity kinematics are lacking.
General trunk fatigue and lower extremities. Trunk fatigue alters the muscle activations
in postural tasks (Hart et al., 2009; Helbostad et al., 2010; Kulas et al., 2010). Hart et al. (2009)
11
investigated isometric lumbar paraspinal muscle fatigue in healthy adults and adults with a
history of low back pain during jogging tasks. After the fatiguing protocol, the healthy control
subjects, on average, increased peak trunk flexion by 0.8°. In comparison to the subjects with
low back pain, who on average decreased peak trunk flexion angle by 0.4°. It is important to
acknowledge that the control group began with an average of 8.9° of trunk flexion prior to the
fatiguing protocol, compared to 10.9° in the low back pain group. Post-fatigue and during
jogging, subjects with low back pain demonstrated greater peak hip abduction angles, 6.6°,
compared to 5.1° in the control group. Along the same lines, subjects with low back pain showed
less peak knee internal rotation during jogging post-fatigue, 10.9° compared to 15.1° in the
control group. Both of these values decreased due to the fatiguing protocol. In comparison to
individuals with low back pain, healthy subjects showed greater forward-flexion of the trunk as
well as less lordotic flexion of the trunk, suggesting trunk fatigue influences lower extremity
kinematics in healthy and unhealthy individuals differently. Specifically, Hart et al. (2009) argue
that individuals with a history of low back pain have developed compensatory muscle activation
patterns to adjust for the increasing fatigue of the paraspinals, exhibiting as fewer postural
adjustments while jogging compared to healthy control subjects.
Lower extremity kinematics are influenced by trunk fatigue (Bellew & Fenter, 2006;
Davidson et al., 2009; Helbostad et al., 2010; Kavanagh et al., 2006; Lin et al., 2009; Petrella et
al., 2005; Pline et al., 2005). After fatigue, balance during voluntary movements is significantly
reduced (Bellew & Fenter, 2006; Helbostad et al., 2010; Helbostad, Leirfall, Moe-Nilssen, &
Sletvold, 2007; Petrella et al., 2005). In an aging population, Bellew & Fenter (2006)
demonstrated that isokinetic fatigue of varying joints significantly diminished balance test
results; fatigue of the lumbar extensors results in increased postural sway during standing tasks
12
(Lin et al., 2009) and increased center of pressure velocities following external perturbations
while standing (Davidson et al., 2009). Trunk fatigue may also decrease lumbar proprioception
and ankle joint position sense (Pline et al., 2005)
General influence of fatigue on the lower extremities. Aerobically induced muscle
fatigue can alter the kinematics of human movement (Dingwell et al., 2008). EMG median
frequency (MDF) was used to track the muscle activation while riding a cycle ergometer at
100% VO2max until voluntary exhaustion. While no direct correlation between specific muscle
MDF and trunk lean occurred, Dingwell et al. (2008) demonstrated that fluctuations in MDF
consistently and significantly preceded fluxes in trunk lean, trunk lean range of motion and hip
angles. The onset of fatigue will result in varying compensatory postures, depending on the
postural and muscular tendencies of the individual (Dingwell et al., 2008; Helbostad et al.,
2010). In many cases, performance deficits occur as a result of multiple factors beyond localized
muscular fatigue including energetic, respiratory and cardiovascular factors (Byrne, Twist, &
Eston, 2004; Dingwell et al., 2008).
Depending on the location of fatigue, fatigue will alter the kinematics of the trunk and
lower extremities differently. For example, lower extremity neuromuscular fatigue increased hip
flexion in men and women and knee flexion in men during a drop jump (Kernozek et al., 2008;
Walsh et al., 2012). Research demonstrates that trunk positioning relative to hips is influential on
knee anterior shear forces (Kulas et al., 2010). However, also as demonstrated in a drop landing,
aerobic fatigue decreased hip and knee flexion, which increases anterior tibial translation and
increases strain on the ACL (Quammen et al., 2012).
Kellis and Kouvelioti (2007) published that consequences of fatigue are highly dependent
on the muscle that was fatigued. Specifically, when knee extensors were fatigued, ground
13
reaction forces significantly decreased during a drop landing, and knee flexion angles were
increased. Specifically, maximal knee flexion and knee flexion at initial contact significantly
increased after extension fatigue (61.46±9.77° and 18.5±4.09° respectively). When knee flexors
were fatigued, maximal knee flexion and knee flexion at initial contact also significantly
increased after fatigue (57.84±11.68° and 15.00±3.74° respectively) (Kellis et al., 2009).
Kinematic risk factors for ACL injury. Multiple factors have been identified to
increase the risk of ACL injury, including knee valgus (Boden et al., 2009; Hewett et al., 2005),
decreased knee flexion (Walsh et al., 2012) increased hip flexion (Boden et al., 2009; Kernozek
et al., 2008) and increased hip internal rotation (McLean, Huang & Bogert, 2005). Boden et al.
conducted a video analysis comparing the mechanics of subjects during ACL injury, compared to
videos of controls performing similar movements. Results demonstrated that peak hip flexion
was greatly increased in subjects compared to controls. Also, knee valgus trended toward
increased values in subjects compared to controls, although results did not reach significant
levels (Boden et al., 2009).
Walsh et al. (2012) identified decreased knee flexion as a risk factor for ACL injury. In a
jump-landing task, Walsh et al. observed sagittal plane lower-extremity kinematics in
recreationally active volunteers during a landing task. Results demonstrated that a more extended
knee, as identified by increased quadriceps activation, was correlated with an increase in impact
forces at the knee, and therefore increased the risk of ACL injury (Walsh et al., 2012). These
findings are corroborated by findings of a previous study conducted by Boden et al. (2000).
Boden et al. interviewed 89 athletes, representing 100 knee injuries, in regards to the
mechanisms about the maneuver of their injury. The majority of the non-contact injuries
occurred at heel strike, when the knee was near full extension (Boden et al., 2000).
14
In a study comparing male and female NCAA athletes sidestep mechanics, McLean et al.
(2005) found a high association between knee valgus and hip internal rotation, specifically in
females. Subjects were required to perform a cutting maneuver between 35° and 55° degrees
from the original direction of motion. Results demonstrated that initial postures with greater
degrees of hip internal rotation, significantly predicted large changes in knee valgus.
Investigators speculate that the increased hip internal rotation inhibited medial muscle groups to
control and limit dynamic knee valgus (McLean et al., 2005).
Lower extremity kinematics. Injured subjects respond to landing tasks either flatfooted
or with the hind foot. This is in comparison to healthy individuals who contact the ground with
their forefoot. The injured subjects, as a consequence of their landing position, showed
significantly greater hip flexion angles (Boden et al., 2009). On average, injured subjects
demonstrated hip flexion angles of 52.4°±17.4° in comparison to control flexion values of
33.4°±17.4° (Boden et al., 2009). Boden et al. did not find a significant difference in hip
abduction angles between injured and healthy controls, although data trended toward lesser
degrees of hip abduction in injured subjects (24.7°±12.6° and 28.4°±10.9° respectively). Foch et
al. (2014) argued that decreased postural control, reflecting as increased lateral flexion of the
trunk contribute to the increased knee adduction observed while running. Interestingly, the
control subjects showed greater hip adduction angles than the iliotibial band syndrome subjects
(15.0°±3.3° and 13.1°±2.6° respectively). These findings are unique in that the control group
showed an increased hip adduction angle, a result typically found in symptomatic individuals
(Foch, 2014).
In line with changes in joint angles during lower extremity tasks, women with histories of
an ACL injury demonstrate lower extremity kinematics that likely predispose them to future
15
injury (Ortiz et al., 2007). Ortiz et al. (2007), published that during a single leg up-down step
test, women with a history of ACL injury executed greater knee valgus angles when compared to
healthy controls (6.09°±5.02° and 5.80°±3.84° respectively). While these findings did not reach
statistical significance, investigators point out that the up-down test, as it is performed primarily
in the sagittal plane, is not a sensitive indicator of lower extremity kinematics in relationship to
ACL injury. In fact, after reconstruction, Ortiz et al. speculate that the experimental subjects
have demonstrated compensatory movement patterns, limiting their degree of knee valgus (Ortiz
et al., 2007).
Lower extremity muscle activation patterns. During normal gait in individuals without
a history of ACL injury, stance support is provided primarily by the gluteus maximus after footflat and before contralateral toe-off (Anderson & Pandy, 2003). Surprisingly, in individuals with
patellofemoral pain, increases in gluteus maximus (GM) activity during a drop jump were
observed (Souza & Powers, 2009). Souza et al. (2009) speculated that the increase in GM
activity during the drop jump was a result of recruiting weakened muscles in attempt to stabilize
the hip structure. Since the GM is a primary abductor, weakness in the GM will increase knee
valgus; a key risk factor for ACL injuries in females (Boden et al., 2009; Hewett et al., 2005).
Additionally, muscular imbalances between the quadriceps and hamstrings in women
have been identified as risk factors for ACL injury. Specifically, excessive activation of the
quadriceps compared to the hamstrings (Q:H) mechanically strains the ACL, related to the forces
exerted on the patella from the quadriceps (Colby et al., 2000; Kernozek et al., 2005; Malinzak,
Colby, Kirkendall, Yu, & Garrett, 2001; Myer, Ford & Hewett, 2004; Myer et al., 2009; Shultz et
al., 2009). Myer et al. (2004) further investigated the over-activation of the quadriceps in females
and confirmed a difference in activation patterns between the medial and lateral quadriceps
16
compared to men. Specifically, a low ratio of medial:lateral quadriceps activation in females was
discussed to increase the strain placed on the ACL by limiting the control of dynamic knee
valgus (Myer et al., 2004).
Cowling & Steele (2000) implemented a rapid deceleration task and measured activations
of the quadriceps and hamstrings comparing men and women. Relative to initial contact during
the rapid deceleration task, males demonstrated a delayed activation of the semimembranosus
compared to females. Investigators discussed that this delay in hamstring activation reflected a
more optimal timing of hamstring activation associated with the high anterior forces just after
initial contact and subsequently relieving some of the strain placed on the ACL. Activation
patterns of the biceps femoris were also reported although they did not reach statistical
significance. By synchronizing the activation of the hamstrings with the peak anterior forces at
the knee, males were able to better accommodate the demand on the ACL (Cowling & Steele,
2000).
Fatiguing protocols. Many protocols have been used in testing and inducing muscular
fatigue of the trunk. Because the trunk is a complex network of musculature of primarily fatigueresistant fibers (Page & Descarreaux, 2012; Richardson et al., 2002), objectifying fatigue can
prove difficult. McGill et al. (1990) even argue that men and women demonstrate varying
endurance performances for trunk endurance tasks. Specifically, females demonstrated greater
trunk extension endurance, while men demonstrated greater lateral and forward flexion
endurance.
Procedures stem from the assumption that fatigue is the inability to maintain a maximal
voluntary contraction (MVC) or a threshold of activation. Objective measurements of MVC
range from peak torque to mean power frequencies of electromyography (EMG) activity.
17
Dedering et al. (2000) investigated the between-days reliability of both subjective and objective
evaluations of back muscle fatigue. Results reflected that EMG frequencies, including mean and
median power frequencies, reflected the subjective reports of fatigue, especially during
prolonged contractions. Reliability of these measurements were greater as contraction times
increased. Moreover, test-retest reliability was increased in active subjects (Dedering et al.,
2000).
In fatigue testing, MVCs are a common procedure for monitoring muscle fatigue. Pretests
MVCs provide a threshold to determine fatigue onset. Depending on the protocols used, the use
of MVC data varies. For example, when MVCs are used to measure fatigue due to an
intervention, pre- and post- test MVC data are compared for changes in muscle activation (Corin
et al., 2005; Page & Descarreaux, 2012). Alternatively, when fatigue is the intervention, MVCs
are used to determine a threshold of activity needed to maintain to induce fatigue (Corin et al.,
2005; Dedering et al., 2000; Ng et al., 2003; Olson, 2009). Common in research, when a
reduction of greater than 10% of the pre-test MVC measurement is reached, fatigue is assumed
(Ament, Bonga, Hof, & Verkerke, 1993; Johnston, Howard, Cawley, & Losse, 1998; Ng et al.,
2003; Jaap H. van Dieën, Selen, & Cholewicki, 2003). Johnston et al. (1998) investigated the
influence of lower extremity fatigue on balance. Using an isokinetic dynamometer, fatigue of the
lower extremity was said to have occurred when force dropped below 50% of the initial testing
muscle activation.
Isometric and isokinetic contractions offer different approaches toward generating
fatigue. Corin et al. (2005) compared the validity of using isometric and isokinetic contractions
for the testing and production of trunk fatigue. Results demonstrated that isometric contractions
were not only a valid assessment of fatigue, but also showed the greatest reduction in peak
18
torque values during the fatiguing protocol. In agreement, multiple studies use isometric
contractions in their fatiguing protocol (Corin et al., 2005; Dedering et al., 2000; De Luca, 1984;
Hart et al., 2009; Ng et al., 2003; Page & Descarreaux, 2012) to generate maximum levels of
fatigue in short timeframes.
Drop landing protocols. Drop landings are used to highlight alterations in kinematics,
muscle activation patterns or movement patterns in the lower extremities (Blackburn & Padua,
2008; Decker, Torry, Wyland, Sterett, & Steadman, 2003; Kellis & Kouvelioti, 2009; Kernozek
et al., 2008; Kulas et al., 2010; Walsh et al., 2012) as deceleration forces demand more from the
musculature and joints (Kellis & Kouvelioti, 2009; Zazulak et al., 2005). Ground reactive forces
during single-leg drop landings can reach 11 times body weight requiring significant muscle
activation to decelerate the body mass (McNitt-Gray, 1993), a benefit of testing lower extremity
kinematics with a drop landing. Along similar lines, non-contact ACL injuries, occurring from
abrupt deceleration or change in direction, are more closely mimicked during drop landings
(Shimokochi, Ambeganokar, Meyer, Lee, & Shultz, 2013)
Killis et al., (2009) implemented single-leg drop landings to determine the effects of
thigh muscle activity on ground reactive forces. Examining gender differences in landing
mechanics, Kernozek et al., (2008) implemented a single-leg, 50cm drop landing procedure with
college athletes. After fatiguing the lower extremities with a squatting protocol, Kernozek and
colleagues observed that women demonstrated significantly more postural alterations than men
did after fatigue during a drop landing. Specifically, women landed with greater hip flexion
(p=0.012) and larger peak valgus overall (p=0.010). Perhaps as a result of inherent anatomical
differences, females demonstrate greater hip adduction and internal rotation, correlating with an
19
increase in knee valgus and external rotation in the knee (Zazulak et al., 2005), a position which
increases risk for ACL injury.
Using a two-foot, 45cm drop landing, Kulas et al., (2010) validated that adding an
external load to the trunk of their participants resulted in increased muscle forces, with the
muscle activation changes being dependent on the trunk orientation upon landing. In a similar
study, Decker et al., (2003) implemented a 60cm drop landing to observe differences in lower
extremity kinematics between genders. Females tended to employ greater hip and ankle joint
range of motions than the males, which increases the energy absorption from the joints closes to
ground impact. Also using a 60cm drop landing, Blackburn and Padua (2008) demonstrated that
an increased trunk flexion upon landing was associated with an increased risk of ACL injury.
Summary
Musculoskeletal fatigue is the inability to perform a given contraction despite sufficient
absolute strength (Allen et al., 2008; Dingwell et al., 2008; Kellis & Kouvelioti, 2009;
MacIntosh et al., 2012; Macintosh & Rassier, 2002; Ng et al., 2003; Sahlin et al., 1998;
Vøllestad, 1997; Walsh et al., 2012; Westerblad et al., 1998). Specifically, peripheral fatigue,
which is a result of a disturbance in substrate concentration within the local muscle environment,
can influence both local and distant muscle fibers (Bellew & Fenter, 2006; Davidson et al., 2009;
Hart et al., 2009; Helbostad et al., 2010; Kavanagh et al., 2006; Kulas et al., 2010; Lin et al.,
2009; Petrella et al., 2005; Pline et al., 2005). Performance discrepancies due to peripheral
fatigue can alter the biomechanics at a joint, reflecting mechanisms of injury at the knee (Boden
et al., 2009; De Luca, 1984; Foch, 2014; Hewett et al., 2005; Horita et al., 1999; Kernozek et al.,
2005; Ng et al., 2003; Shultz et al., 2009).
20
Kinematic measures like increased knee valgus, decreased knee flexion, increased hip
flexion and increased hip adduction reflect increased risk for ACL injury (Boden, Torg,
Knowles, & Hewett, 2009; Hewett et al., 2005; Souza & Powers, 2009). Additionally, an
increase in Q:H activation ratios and decrease in gluteus maximus activation are muscle
activation patterns consistent with ACL mechanisms of injury (Boden et al., 2009; Hewett et al.,
2005; Zazulak et al., 2007). This study aims to understand how abdominal fatigue may effect
these biomechanical factors related to ACL injury.
21
Chapter III
Methodology
Introduction
This study tested the hypothesis that abdominal fatigue would alter both hip and knee
kinematics and muscle activation patterns during a drop landing. Fatigue was induced by
maintaining 40° of trunk flexion relative to the ground as measured by an adjustable wedge
created for this study. Fatigue was said to have been achieved when subjects could no longer
maintain the initial 40° posture as determined by the wedge, and their back lowered to make
contact with the wedge at 20° relative to the ground. Three-dimensional motion capture was used
to measure hip kinematics in the frontal, and transverse planes as well as knee kinematics in the
frontal and sagittal plane. Electromyography data was collected from the superior rectus
abdominis (SA), inferior rectus abdominis (IA), external obliques (EO), rectus femoris (RF),
biceps femoris (BF) and gluteus maximus (GM) of the participant’s dominant side during preand post-test drop landings.
Description of study population
The sample used in this study consisted of 27 female college students, between the ages
of eighteen and thirty 21.52 ± 1.16, height 1.62 ± 0.08 m, mass 61.66 ± 12.52 kg. Exclusion
criteria included history of lower extremity or back injury in the last 6 months. As the results of
this study were applied to an average, generally healthy, college aged, female, inclusion criteria
regarding fitness levels and health were null, hoping to include typically healthy young adults.
22
Design of study
The study was a within-subject, repeated measure design to analyze the effect of
abdominal fatigue on hip and knee kinematics. 27, female, college-aged subjects were recruited
from Kinesiology courses at Western Washington University. Subject’s performed a pre- and
post-fatigue test drop landing protocol, demonstrating the effects of abdominal fatigue.
Data collection procedures
The Human Subjects Committee at Western Washington University approved this study.
The potential risks and benefits were described in detail to the subjects, and each subject
completed a written informed consent (Appendix A) prior to the collection of any data. Subjects
were then measured for demographic data, including weight and height (Appendix B).
Instrumentation. The fatiguing protocol utilized surface electromyography (EMG) to
measure muscle activation amplitudes of the external obliques, superior and inferior rectus
abdominis, rectus femoris, biceps femoris and gluteus maximus of the right side. Additionally,
Qualisys three-dimensional motion capture analysis was used to measure the joint kinematics of
the hip and knee during the drop landing.
Surface electromyography. The Noraxon TeleMyo desktop direct transmission EMG
system (Noraxon, Scottsdale, AZ, USA) was used to monitor amplitude. Prior to placing the
electrodes on the subjects, the skin was shaved and cleaned with alcohol wipes to reduce any
noise. Disposable, Noraxon, self-adhesive Ag/AgCl snap electrodes were placed parallel to the
muscle fibers on the external obliques (EO), superior and inferior rectus abdominis (SA and IA
respectively), rectus femoris (RF), biceps femoris (BF) and gluteus maximus (GM) (Figure 2a, b,
c) of the right side (S. McGill, Juker, & Kropf, 1996; Page & Descarreaux, 2012). The interelectrode distance was 11 mm (Figure 1). Muscle bellies of these muscles were palpated by a
23
single investigator using sub-maximal isometric contractions in accordance to Olson et al. (2009)
protocol.
Figure 1: Noraxon Dual Electrodes, inter-electrode distance is 17 mm.
Three dimensional motion analysis. Qualysis motion analysis system (Qualisys Track
Manager, Gothenburg, Sweden) using 7 ProReflex cameras was used to collect kinematic data of
the hip and knee. Specifically, knee joint angles in the frontal and sagittal plane, as well as hip
joint angles in the frontal and transverse plane. Retroreflective markers were attached to bilateral
anterior superior iliac spines, bilateral posterior superior iliac spines, bilateral greater trochanters
of the femur, medial and lateral femoral epicondyles of the right side, lateral and medial malleoli
of the right side, , base of the first metatarsal. Additionally, clusters of four markers were placed
on the lateral aspect of the right thigh and right shank in a rectangular arrangement, mid-distance
from the proximal and distal segment markers. A cluster of three markers were placed on the
foot as well, one lateral to the head of the first metatarsal, and two located midway between these
markers an the maleoli (Hart et al., 2009; Kernozek et al., 2008, 2005; Ortiz et al., 2007) (Figure
2a, b, c).
24
Figure 2a: lateral view of the retroreflective marker placement of the right lower
extremity, as well as the electrical goniometer can be seen in alignment with the right
femur and tibia. Cluster markers are placed on the lateral aspects of the thigh and shank,
as well as on the dorsal aspect of the foot. Joint landmark markers can be seen on the iliac
crest, posterior superior iliac spine, anterior superior iliac spine, greater trochanter of the
femur, lateral epicondyle of the femur and lateral malleolus.
Figure 2b: posterior view of the retroreflective marker placement of the pelvis and right
lower extremity. Medial epicondyle and malleolus markers are visible. EMG electrodes
of the BF can also be seen.
25
Figure 2c: anterior view of the retroreflective marker placement of the pelvis and SA, IA
and EO EMG electrodes. Iliac crest, anterior superior iliac spines as well as greater
trochanter markers can be seen.
Measurement techniques and testing procedures. Data collection occurred at Western
Washington University within the Biomechanics lab in Carver Gymnasium. Upon entry, each
subject reviewed the informed consent form (Appendix A), was offered an opportunity to ask
questions, and then was asked to sign the consent form. Additionally, subject’s height and weight
was recorded and were questioned regarding injuries to their lower extremities or trunk within
the last 6 months (Appendix B). Subjects that affirmed a recent history of injury were excluded
from the study.
In agreement with relevant research, participants were required to perform a 5 minute,
self-selected intensity cycling warm up prior to instrumentation (Blackburn & Padua, 2008;
Kellis & Kouvelioti, 2009; Kernozek et al., 2005; Walsh et al., 2012). Upon completion of the
warm up, a single investigator demonstrated one drop landing, with minimal verbal instruction to
avoid influencing the subject’s natural movement pattern. Each subject was instructed in the
same manner as to how to perform a drop landing : “on the command of ‘Go’, step off the box
with your right foot, land comfortably and naturally with both feet simultaneously, stand up
straight after landing, and then step backwards on to the box with your left foot”. Subjects were
26
then allowed to practice drop landings until comfortable with the procedure with minimal input
to ensure natural landing mechanics.
Subjects were then instrumented with EMG electrodes on the right SA and IA, and right
EO (Dedering et al., 2000), as well as the RF, BF and GM of the right leg (Table 1). An
electronic goniometer was placed about the knee, with the proximal portion in line with the
femur, and the distal portion in line with the tibia (Figure 2a). A single standing trial was
collected for 3 seconds with the Noraxon computer system to identify each subject’s preferred
stance. The participant then performed a single voluntary isometric reference contraction for
each muscle group (Table 1). In order to normalize the EMG amplitude, reference contractions
were performed using a maximum voluntary isometric contraction (MVIC). Reference
contractions allow data to be presented as a percentage of activation, further allowing data to be
compared between subjects relative to their own capacity for contraction.
Table 1. Electrode Placement (Olson et al., 2010) and Reference Contractions
Muscle Group
Superior
Rectus
Abdominis
Palpation technique and
electrode placement
A trunk extension force was applied
at the shoulders while subjects were
standing. Subjects were instructed
to resist this force and provide a
trunk flexion action. The muscle
belly of the superior rectus
abdominis was palpated just lateral
to the midline of the body and distal
to the last intercostal space.
Electrodes were placed parallel to
the muscle fibers.
27
Reference contraction technique
Subjects were supported by the
wedge at 40°, knees at 90° and feet
secured with a 30 pound sandbag.
Subjects were instructed to support
their head against the wedge, arms
across their chest. A research
assistant placed their hands on their
shoulders, placing a trunk-extension
force on the subject, holding them
into the wedge. Subjects were
instructed to perform a trunk flexion
action against the research assistant
maximally. Research assistants were
instructed to match the subject’s
resistance to maintain an isometric
contraction.
Palpation technique and
electrode placement
Inferior Rectus A trunk extension force was applied
at the shoulders while subjects were
Abdominis
standing. Subjects were instructed
to resist this force and perform a
trunk flexion action. The muscle
belly of the inferior rectus
abdominis was palpated just lateral
to the midline of the body, distal to
the umbilicus. Electrodes were
placed parallel to the muscle fibers.
A lateral trunk flexion force was
External
applied at the right shoulder
Oblique
(subject was pulled to their left).
Subjects were instructed to resist
this force and provide a trunk
flexion action in the opposing
direction (flexion to their right).
The muscle belly of the external
oblique was palpated parallel to the
line made by the iliac crest and the
anterior superior iliac spine.
Electrodes were placed parallel to
the muscle fibers.
Subjects were seated, knees
Rectus
supported at 90° flexion. A flexion
Femoris
force was applied to the shank.
Subjects were instructed to resist
this force by producing a knee
extension action. The muscle belly
of the rectus femoris was palpated
in the superior portion of the thigh.
Electrodes were placed parallel to
the muscle fibers.
Muscle Group
28
Reference contraction technique
Reference contraction was performed
in the same manner as the superior
rectus abdominis and at the same
time.
Reference contraction was performed
in the same manner as the superior
rectus abdominis and at the same
time.
Subjects were seated, thighs
supported, and shanks unsupported,
perpendicular to the ground. A knee
flexion force was applied at the
shank, just proximal to the malleoli.
Subjects were instructed to perform a
maximal knee-extension action.
Research assistants matched the
maximal force applied by the subject
to maintain an isometric contraction.
Palpation technique and
electrode placement
Biceps Femoris Subjects stood, facing into a table
and were instructed to lift their
right shank to 90° relative to their
thigh. A knee extension force was
applied at the distal end of the
shank. Subjects were instructed to
resist this force by performing a
knee flexion action. The muscle
belly of the biceps femoris was
palpated in the superior and lateral
aspect of the posterior thigh and
electrodes were placed parallel to
the muscle fibers.
Subjects stood facing away from
Gluteus
the seated investigator, and placed
Maximus
their right foot on the investigator’s
shank. Subjects stood at
approximately 35° of hip flexion
and were instructed to press
maximally into the shank of the
investigator with their heel. The
muscle belly of the gluteus
maximus was located by
identifying the line created by the
right greater trochanter and the
right posterior superior iliac spine,
at the peak of the muscle.
Electrodes were placed parallel to
the muscle fibers.
Muscle Group
Reference contraction technique
Subjects stood, facing into a table at
hip height, hands resting on the table.
Subjects were instructed to flex their
knee so that their shank was parallel
to the floor. A research assistant
applied a knee-extension force, while
subjects were instructed to perform a
maximal knee-flexion action.
Research assistants matched the
maximal force applied by the subject
to maintain an isometric contraction.
Subjects stood with their back to a
wall, and a chair placed in front pf
them for stabilization. Subjects were
instructed to flex their knee so that
their shank was parallel to the
ground, and stand at a distance
between the wall and chair so that
their hip was approximately at 35° of
flexion. Subjects were then instructed
to press into the wall maximally with
their heel.
All reference contractions were performed in the following order: RF, BF, GM, and
abdominals (superior and inferior rectus abdomonis, and external oblique). Following the
reference contractions, subjects were allowed two minutes rest (Kernozek et al., 2005), during
which time they were instrumented with retroreflective markers on bilateral anterior superior
iliac spines, posterior superior iliac spines, greater trochanters of the femur, medial and lateral
femoral epicondyles, a cluster on the lateral aspect of the thigh, a cluster on the lateral aspect of
29
the shank, lateral malleoli, medial malleoli, base of the first metatarsal, and a cluster on the foot
for three-dimensional motion capture as previously stated in accordance with recent research
(Hart et al., 2009; Kernozek et al., 2008, 2005; Ortiz et al., 2007).
Once fully instrumented with both surface EMG electrodes and the retroreflective
markers, subjects performed three consecutive drop landings (Kellis & Kouvelioti, 2009). Drop
landings were performed from a box with a height of 30.5 cm (Blackburn & Padua, 2008;
Decker et al., 2003; Kellis & Kouvelioti, 2009; Kernozek et al., 2005; Shultz et al., 2009; Walsh
et al., 2012) (Figure 3). Subjects were counted-down for the first drop landing, and instructed to
self-pace for the remaining two drop landings while taking minimal time between trials.
Kinematic data from the first and second trials were used for further analysis. The third trial was
used in case of poor data collection during the other trials.
Adjustable Wedge at 40 ̊
Drop Landing Box
Figure 3: The drop landings were performed from a height of 30.5cm. The adjustable
wedge for the fatiguing protocol was placed within the capture area, so that the time from
fatigue to post-testing could be minimized.
30
Fatiguing occurred in the data collection area to allow for minimal transfer time between
fatigue and post-testing (Figure 3). After pre-test data collection, the fatiguing protocol
commenced. Subjects were asked to maintain their torso at a 40° angle relative to the ground
with assistance from the wedge (Figure 4), knees flexed to 90°, and feet secured. Once oriented
in the proper posture and joint angles, the wedge was removed away from the subjects 2 inches
(McGill, 1999), and lowered to 20° (Figure 5). Subjects were instructed to maintain the initial
40° posture “for as long as and as accurately as possible”. Abdominal fatigue was said to have
been achieved when subjects could no longer maintain the 40° posture and their back lowered to
make contact with the wedge again. Subjects were given 5 seconds to move from the fatiguing
location to the post-test collection.
Figure 4: Subjects began supported by the wedge, with their arms across their chest,
knees flexed to 90̊ as confirmed by the electrical goniometer, and feet secured with a sand
bag.
31
Figure 5: Subjects maintained the initial 40̊ position for as long and as accurately as
possible, while the wedge was lowered 20 ̊ and removed back by 2 inches.
Immediately after reaching fatigue, subjects performed three additional consecutive drop
landings (Kellis & Kouvelioti, 2009). Subjects were instructed to transition between droplanding trials quickly, in the same manner as in the pre-test. Kinematic data from the first and
second trials were used for further analysis. Upon completion of the testing procedures, subjects
were de-instrumented and allowed time to recover from the fatiguing protocol (Blackburn &
Padua, 2008; Kellis & Kouvelioti, 2009; Kernozek et al., 2005; Walsh et al., 2012).
Data Processing.
Pre- and post-test droplandings were cropped in Qualisys Track Manager(Qualisys Track
Manager, Gothenburg, Sweden) as to separate drop landings. Individual drop landings were
identified as the movement occurring between the most extended position at right knee after
losing contact with the drop landing box, and the subject standing erect after landing. Peak joint
angles were assumed to have occurred while the feet were in contact with the ground. Graphing
of the goniometer data was used to identify the onset of knee flexion as the first inflection point
in the positive flexion direction, and the point where knee flexion returned to neutral or extension
32
was identified as the completion of the landing. A t-test was used to compare the pre- and posttest muscle activation profiles of the first landings. SPSS (Version 22, Chicago, IL, USA) was
used to complete the statistical analysis.
Data Analysis.
Data analysis of the joint kinematics was performed using a 1-way ANOVA, comparing
the pre-test and post-test data for frontal and sagittal plane knee kinematics, as well as frontal
and transverse plane hip kinematics. These measures included knee flexion, knee valgus, hip
flexion and hip internal rotation. Peak angles during the second pre-test landing, and the first and
second post-test landings were compared using repeated measures 1-way ANOVA to compare
the effects of time from pretest to post test, as well as first drop landing to second drop landing.
Maximum voluntary isometric contraction (MVIC) was used to normalize the muscle activation
data so that results could be compared between subjects. Muscle activation occurring between
initial knee flexion completion of the landing, as determined by the goniometer data in MR3
(Noraxon, Scottsdale, AZ, USA) was averaged and analyzed for peak and mean activation.
33
Chapter IV
Results and Discussion
Introduction
This study tested the hypothesis that abdominal fatigue would result in altered lowerextremity kinematics and muscle activation patterns, consistent with risk factors of ACL injury.
Specifically, the experimental hypothesis was that abdominal fatigue would cause a decrease in
knee flexion, increase in knee valgus, increase in hip flexion and increase in hip internal rotation.
In regards to muscle activation, an increase in quadriceps to hamstring ratio (Q:H) by an increase
in quadriceps activation and a decrease in hamstring activation, and decreased gluteus maximum
activation were hypothesized. Peak joint angles of the knee and hip were measured during three
drop-landings pre- and post- abdominal fatigue. Additionally, muscle activation of the RF, BF,
GM, SA, IA and EO of the right side was collected. Post-testing occurred 5-seconds post-fatigue.
A repeated measures 1-way ANOVA analysis was conducted to compare the joint kinematics of
the second pre-test landing to the first and second post-test landings. Additionally, a paired
sample t-test was used to compare muscle activation profiles of the first pre-test and post-test
landings.
Results
Time to fatigue across all subjects ranged from 49 seconds to 795 seconds (176 ± 6.74).
There was a significant effect of time on knee flexion (F[2, 47.323]=6.194, p=0.004).
Additionally, peak hip flexion and hip internal rotation increased between pre-test and post-test
measures, although there was not a significant effect of time (p>0.05). In regards to muscle
activation, activation of the rectus femoris was significantly increased from pre-test to post-test 2
34
(p<0.05), as well as a non-significant decrease in gluteus maximus activity (p>0.05). Raw data is
provided in Appendix C.
Joint Kinematics. Joint kinematic data were analyzed for peak joint motion in a given
plane between initiation of the drop landing to the completion of the landing, or the most erect
posture achieved after landing. These data stems from the assumption that peak joint angles
occurred while the subject was in contact with the ground. Based on Mauchly’s test for
sphericity, maximum knee flexion data met the assumption of sphericity for the condition of
time. Therefore, sphericity was assumed. There was a significant effect of time on maximum
knee flexion (F[2, 47.323]=6.194, p=0.004). A Bonferoni correction was applied, and post-hoc
testing revealed a significant increase from pre-test to both post-test 1 and post-test 2 measures
(p=0.012 and p=0.006 respectively) (Figure 6). Mauchly’s test revealed that data for maximum
knee valgus did not violate the assumption of sphericity for the condition of time and sphericity
was assumed. There was not a significant effect of time on maximum knee valgus (F[2,
52]=1.197, p=0.728). However, maximum knee valgus increased post-test from pre-test (10.66°),
and was greater in the second post-test (11.04°) landing than the first (10.99°) (Figure 7).
35
Maximum Knee Flexion
94
*
*
Post-1
Post-2
Knee Flexion (Degrees)
92
90
88
86
84
82
0
80
Pre test
Figure 6. A graphical comparison of maximum knee flexion during pre-test, and the first and
second post test landings. * p<0.05
Mean Knee Valgus
14
13
Knee Valgus (Degrees)
12
11
10
9
8
7
6
50
Pre test
Post-1
Post-2
Figure 7. A graphical comparison of maximum knee valgus across subjects between pre-test and
the first and second post-test landings. A less positive number represents a greater degree of knee
valgus.
36
Mauchly’s test revealed that the data for maximal hip flexion violated the assumption of
sphericity for the condition of time. Therefore, the Greenhouse-Geisser correction for degrees of
freedom was applied for this effect. No significant effect of time was found between pre-test and
post-test measures of hip flexion (F[1.126, 28.143]=2.095, p=0.157). Peak hip flexion increased
from pre-test to post-test 2 (67.98° and 71.42° respectively), although not to a significant extent
(Figure 8). With the hip internal rotation data violating the assumption of sphericity, the
Greenhouse-Geisser correction for degrees of freedom was applied. No significant effect was
found between pre-test and post-test measures of hip internal rotation (F[1.558, 1.269]=38.951,
p=0.932). Internal hip rotation did increase pre-test to post-test, with the largest increase
occurring between the first and second post-test landings (-11.23° and -10.91° respectively),
although these results did not reach statistically significant amounts (Figure 9).
Maximum Hip Flexion
75
Hip Flexion (Degrees)
70
65
60
55
50
45
40
35
0
30
Pre test
Post-1
Post-2
Figure 8. A graphical comparison of the changes in maximum hip flexion between pretest, and post-test landings.
37
Maximum Hip Internal Rotation
Hip Internal Rotation (Degrees)
10
9
8
7
6
5
4
3
2
1
0
Pre test
Post-1
Post-2
Figure 9. A graphical comparison of the mean internal rotation of the hip between pretest
and post-test landings. A less positive number reflects a greater degree of internal
rotation.
Muscle Activation. Mean muscle activation for the abdominal muscles varied; post-test,
SA activation decreased, IA activation increased slightly, and EO activation increased greatly,
although statistical significance was not achieved (p>0.05). Similar patterns were observed in
peak muscle activation of the SA, IA and EO (p>0.05). Mean activation for the RF
insignificantly increased post-test (Figure 10), while peak activation of the RF significantly
increased post-fatigue (p<0.05) (Figure 11 & 12). Activation of the BF decreased in both mean
activation and peak activation, although statistical significance was not achieved (p>0.05)
(Figure 10 & 11). Additionally, GM activation decreased in both mean activation and peak
activation, although not reaching statistical significance (p>0.05) (Figure 10 & 11). Finally,
although not significant, peak Q:H increased from 3.07 ± 0.52 to 4.03 ± 0.85 after the fatiguing
38
protocol (Figure 13 & 15). Similarly, mean Q:H ratio increased from 1.38 ± 0.07 to 3.19 ± 0.39
after fatigue (Figure 14 & 16).
Mean Muscle Activation
Pre Test
Post Test
30
25
MVC (%)
20
15
10
5
0
RF
GM
BF
Figure 10. A graphical comparison of mean muscle activation for the RF, GM and BF
pre-test and post-test. Standard error bars are represent the mean standard error of the
means.
Peak Muscle Activation
Pre Test
Post Test
70
60
*
MVC (%)
50
40
30
20
10
0
RF
GM
BF
Figure 11. A graphical comparison of average peak muscle activation for the RF, GM,
and BF pre-test and post-test. * p<0.05
39
Peak Quadriceps Activation
Pre Test
Post Test
70
*
Normalized Activation (%)
60
50
40
30
20
10
0
Figure 12. Comparison of mean peak quadriceps activation pre-test and post-test.
*p<0.05
Mean Quadriceps to Hamstring Comparison
Rectus Femoris
Biceps Femoris
Normalized Activtation (%)
35
30
25
20
15
10
5
0
Pre Test
Post Test
Figure 13. A graphical representation of the mean activation differences between the RF
and BF pre- and post-test.
40
Peak Quadriceps to Hamtring Comparison
Rectus Femoris
Biceps Femoris
70
60
MVC (%)
50
40
30
20
10
0
Pretest
Posttest
Figure 14. A graphical representation of the peak activation differences between the RF
and BF pre- and post-test.
Peak Q:H Ratio
12
10
8
6
4
2
0
Pretest
Posttest
Figure 15. A graphical representation of the peak Q:H activation ratios pre- and postfatigue.
41
Mean Q:H Ratio
8
7
6
5
4
3
2
1
0
Pretest
Posttest
Figure 16. A graphical representation of the mean Q:H activation ratios pre- and postfatigue.
Discussion
The purpose of this study was to identify if acute abdominal fatigue would increase risk
factors for ACL injury, specifically those regarding lower extremity kinematics and muscle
activation patterns. The results support the hypothesis demonstrating that abdominal fatigue
significantly increases specific risk factors of ACL injury in healthy, college-aged females.
Kinematic results from this study are limited in that they are based on the assumption that peak
joint angles occurred while the subject was in contact with the ground, not during the flight
phase. However, results demonstrated that maximal knee flexion significantly increased from
pre-test, and maximal hip flexion increased, although not significantly. Maximum knee valgus
increased between pre-test and post-test although not significantly. Results also showed an
increase in hip internal rotation, with maximal values demonstrating a non-significant increase
from pre-test to post-test 2 measurements.
42
Muscle activation pattern results showed a significant increase in RF activation post-test
compared to pre-test (18.62% and 27.06% respectively, p=0.015). Additionally, there was a
decrease in mean BF activation (14.07% and 9.88%, pre-test and post-test respectively), however
these differences were not statistically significant. As a result of these two changes, peak Q:H
ratio increased from 3.07 ± 0.52 to 4.03 ± 0.85 after fatigue, and mean Q:H ratio increased from
1.38 ± 0.07 to 3.19 ± 0.39 . While not significant, both mean and peak GM activation decreased
post-fatigue.
Mechanisms of increased risk for ACL injury include a decrease in knee flexion during a
landing (Walsh et al., 2012). In disagreement, results from this study demonstrated significant
increases in maximum knee flexion angles. This may be due to the RF fatiguing as a hip-flexor
during the abdominal fatigue protocol, and therefore not being able to act eccentrically to
decelerate through the drop-landing, increasing knee flexion. Literature states that the decrease in
knee flexion is associated with an increase in quadriceps activation, a primary knee extender
(Walsh et al., 2012), although despite the current study’s significant increase in RF activity,
results demonstrated increases in knee flexion.
Results from this study demonstrated significant increases in maximum knee flexion
angles (p<0.05). Numerous studies have identified decreases in knee flexion as a risk factor for
ACL injury by comparing injury-maneuvers to maneuvers that did not cause injury (Boden et al.,
2009 & Hewett et al., 2005). Similarly, Walsh et al. (2012) identified a decrease in knee flexion
as a risk factor for ACL injury after making the assumption that an increase in quadriceps
activity would reflect a more extended knee. In this case, Walsh et al. (2012) correlated the
quadriceps activation with a decrease in knee extension. In the current study, if the RF was
fatigued as a hip-flexor during the abdominal fatigue protocol, this could limit the ability of the
43
RF to eccentrically decelerate during the landing, therefore increasing knee flexion. In
corroboration with these results, Kellis & Kouvelioti (2009) observed that knee extensor fatigue
caused an increase in knee flexion. Again, under the assumption in the current study that the RF
was fatigued during the abdominal fatigue protocol, and with the understanding that the RF acts
both as a hip flexor and a knee extensor, results are consistent with those of Kellis & Kouvelioti
(2009). In comparison, Kellis & Kouvelioti (2009) reported maximum knee flexion angles
during a drop landing to be 61.46° ± 9.77° after knee extensor fatigue, while maximum knee
flexion angles of this study reached 89.98° ± 13.22°. These differences in maximum knee flexion
are likely due to the differences in fatigue protocols, but may also be due to differences in the
demographics of subjects. Specifically, Kellis & Kouvelioti (2009) recruited slightly older, taller,
and lighter subjects than the subjects of the current study.
The present study did not observe any significant changes in knee valgus after abdominal
fatigue (p>0.05). Ortiz et al. (2007) performed an up-down task, looking to identify risk factors
for ACL injury. After not finding significant differences in individuals with or without ACL
injury, investigators identified that the specific task was performed primarily in the frontal plane,
and would not be sensitive to sagittal plane movements of the knee. This assumption could be
directed toward the current study as well, explaining why no significant differences in knee
valgus were observed. Knee valgus is consistently identified as a risk factor for ACL injury
(Hewett et al., 2005; Kernozek et al., 2005 & 2008). It is also consistently reported that during
drop landings, females land with consistently greater knee valgus than males (Kernozek et al.,
2005 & 2008). Kernozek et al. (2005) identified that fatigue did not alter knee kinematics in the
sagittal plane. Specifically, females landed with 3.86° and 3.91°of knee valgus, pre-fatigue vs.
post-fatigue respectively, results which were not significantly different from each other.
44
Similarly, this study reported no change in knee valgus pre- and post-fatigue. However, results of
this study did demonstrate much greater degrees of knee valgus, reaching 10.66°± 5.07° and
11.04° ± 5.66° pre- and post-test 2 respectively. While both studies reported no change in knee
valgus after fatigue, the current study reported much greater values in comparison to Kernozek et
al. (2005). These differences in amounts of knee valgus may be as a result of differing fatigue
protocols (lower extremity vs. abdominal fatigue), as well as differences in drop landing heights
(50 cm and 30.5 cm, Kernozek et al., 2005 and the current study respectively). Finally, Kernozek
et al. (2005) recruited female athletes compared to generally healthy females. It is possible that
female athletes perform drop landings with different kinematics due to their training status or
training background.
In the current study, maximal hip flexion did not change after fatigue (p>0.05). This is
likely due to the wide variance in preferred landing mechanics across subjects. Literature reports
that an increase in hip flexion increase the risk for ACL injury. McLean et al. (2005) observed a
significant correlation between hip flexion at initial contact of a cutting maneuver and knee
valgus. Specifically, a greater degree of hip flexion at initial contact reflected a greater degree of
knee valgus. McLean et al. (2005) argued that the observed increase in hip flexion may have
limited the ability of the thigh musculature to support knee-valgus loads, explaining the positive
relationship between the joint kinematics. While McLean et al. (2005) did not explicitly report
hip flexion values, based on the graphical representations of the data, it appears that a much
greater degree of hip flexion was observed in the current study. This is likely due to differences
in tasks, with a cutting task requiring less vertical displacement of the center of mass, and a
greater horizontal displacement. In another study, after lower extremity fatigue, Kernozek et al.
(2008) observed a greater than 8º increase in hip flexion. In comparison, the current study only
45
observed an approximately 4º increase in hip flexion after abdominal fatigue. Kernozek et al.
(2008) implemented a 50 cm drop landing, which may offer a further explanation as to why a
greater increase in hip flexion was observed.
While this study did not detect a significant difference in hip internal rotation pre- and
post-test, an increasing pattern was observed. Recent literature identifies increases in hip internal
rotation a risk factor for ACL injury (Frank et al., 2013; Lephart et al., 2002; Sigward et al.,
2014). Frank et al. (2013) identified a relationship between hip internal rotation and an increased
internal knee external rotation; a load associated with increased risk for ACL injury. Similarly,
Sigward et al. (2014) found that during a cutting maneuver, females performed with significantly
greater degrees of hip internal roation. In fact, in female subjects, Sigward et al. (2014) found
that hip internal rotation explained 63% of the variance observed in knee valgus. Similar to
McLean et al. (2005), investigators argued that increases in hip internal rotation may limit the
ability of the thigh musculature to support valgus loads at the knee (Sigward et al., 2014).
Results of the current study may not have reflected significant differences in hip internal rotation
due to the fact that a vertically based task was implemented, in comparison to a cutting task.
Interestingly, subjects in the current study performed the drop landing task with values of hip
internal rotation similar to individuals with patellofemoral pain (Souza et al., 2009). Comparing
healthy individuals to individuals with patellofemoral pain, Souza et al. (2009) observed subjects
landing with 1.2 ° ± 3.8° and 7.6° ± 7.0° degrees of hip internal rotation respectively. Subjects in
the current study performed a drop landing task with 5.65°± 5.99° and 7.68°± 5.63° pre-test and
post-test 2 respectively. Despite being a healthy population, individuals in the current study
performed with hip internal rotation consistent with those of individuals with patellofemoral
46
pain. It should be identified that pre-test values of hip internal rotation were greater than those of
the reported healthy subjects in the study performed by Souza et al. (2009).
The current study did not observe changes in the ratio of activation between the
quadriceps and hamstrings. However, this study was limited in that only the rectus femoris (one
of four quadriceps muscles) and the biceps femoris (one of three hamstring muscles) were
analyzed. Consequently, an incomplete picture of muscle activation ratios were collected,
perhaps explaining why no significant differences were observed pre- and post-test. Another
factor which may explain why no significant changes were observed could be due to the
previously discussed large variance in landing patterns across subjects. While on average the
data reflects a quadriceps-dominant landing patterns, not all subjects landed in such a manner.
Along similar lines, there was a large standard error of the mean in BF activation data,
contributing to the lack of statistical significance observed in the Q:H ratio. Additionally, if it is
assumed that the RF was recruited and fatigued during the abdominal fatigue protocol, the
increase in RF activation may be due to the fatigue (Ng et al., 2003). Imbalances between the
quadriceps and hamstrings, specifically excessive activation of the quadriceps compared to the
hamstrings (Q:H) has been identified to mechanically strain the ACL, related to the forces
exerted on the patella from the quadriceps (Colby et al., 2000; Kernozek et al., 2005; Malinzak,
Colby, Kirkendall, Yu, & Garrett, 2001; Myer, Ford & Hewett, 2004; Myer et al., 2009; Shultz et
al., 2009). Although, it should be noted that while there were no statistically significant changes
in mean and peak Q:H ratios, both mean and peak ratios increased after fatigue. In fact, the peak
Q:H ratio increased from 3.86 ± 2.72 to 5.67 ± 5.74 after abdominal fatigue, increasing by nearly
1.5 times. Similarly, mean Q:H ratio increased from 3.07 ± 2.09 to 4.03 ± 3.42, increasing 1.25
times. These values are greater than activation ratios reported by Kellis & Kouvelioti (2009)
47
during loading response 1 (100-ms immediately after initial ground contact) possibly due to
differences in fatiguing protocols and subject demographics.
The current study did observe significant increase in RF activation after abdominal
fatigue (p<0.05). As previously mentioned, this could be due to the fact that raw EMG activation
increases with fatigue, and under the assumption that the RF was fatigued during the fatigue
protocol, could partially explain the significant increases observed. The combination of results in
this study including increases in quadriceps activation (p<0.05), increases in knee flexion
(p<0.05) and sustained hip flexion reflect a quadriceps dominant landing pattern. Assuming a
more quadriceps dominant landing pattern after fatigue, is consistent with the significant
increases in RF activation. With that said, due to the forces exerted by the quadriceps on the
patella, increases in quadriceps activation are associated with an increase in ACL injury
(Cowling & Steele, 2000; Kernozek et al., 2008; Ortiz et al., 2008; Shultz et al., 2009).
Specifically, Shultz et al. (2009) identified that subjects, regardless of gender or fatigue, who
landed with less hip flexion, greater knee flexion and increased quadriceps activation all
sustained greater anterior tibial shear forces during the deceleration phase of the landing. These
variables are consistent with those observed in the current study, suggesting that subjects
achieved greater anterior tibial shear forces, a factor associated with increased risk for ACL
injury (Cowling & Steele, 2000; Kernozek et al., 2008; Ortiz et al., 2008; Shultz et al., 2009).
The present study did not observe a significant change in hamstring activation,
specifically BF activation, after abdominal fatigue. However, peak and mean BF activity did tend
to decrease after fatigue. Statistically significant differences were likely not observed due to
large standard errors of the mean in BF activation data. As the quadriceps antagonist, a decrease
in hamstring activity is identified as a risk factor for ACL injury (Cowling, 2000; Kellis, 2009;
48
Kulas, 2010; Shultz, 2009; Walsh, 2012). Cowling & Steele observed that timing of the BF
activation during a drop landing was important in attenuating anterior shear forces at the tibia.
Investigating knee extensor fatigue, Kellis & Kouvelioti (2009) found a significant decrease in
BF activation just prior to landing after fatigue (21.3 ± 4.8% and 16.9 ±5.5% before and after
fatigue respectively). These values are similar to the peak BF activation demonstrated in the
current study (26.5 ± 11.69% and 17.4 ± 4.06% before and after fatigue respectively).
Results of the current study did not demonstrate significant changes in GM activity after
fatigue. As the GM is a hip abductor, weakness in the GM may increase knee valgus; a key risk
factor for ACL injuries in females (Boden et al., 2009; Hewett et al., 2005). In agreement with
the results of this study, Walsh et al. (2012) observed that the GM did not significantly contribute
to support during the deceleration phase of a drop landing, perhaps explaining why no
differences were observed pre- and post-test. Walsh et al. (2012) also identified that an increase
in GM activation during the preparatory phase of the landing (just prior to initial contact) was
correlated with a more extended knee during landing. With this in mind, the observed decrease in
GM activity (p>0.05) may explain some of the increases in knee flexion (p<0.05) observed in
this study.
A number of factors limited this study. For example, the abdominal fatiguing protocol did
not isolate the anterior core musculature, and likely recruited the hip flexor musculature to
maintain the specified posture. By producing hip flexor fatigue, lower extremity kinematics were
likely effected differently than if the abdominals were fatigued in isolation, perhaps explaining
the non-significant increases in hip flexion. However, because it is not likely abdominal fatigue
occurs in isolation, results of this study are more widely generalizable. Additionally, there was a
wide range in time-to-fatigue, ranging from 49 seconds to over 13 minutes, likely due to not
49
controlling for activity level or training status. While this allows broader conclusions to be drawn
by recruiting a generally healthy population, it is possible that trained and untrained individuals
respond to abdominal fatigue differently, thus limiting results. Along similar lines, the wide
range in time-to-fatigue may have been due to a lack of control for the sustained posture during
the fatigue protocol. While subjects were encouraged to maintain the original 40º posture for as
long and as accurately as possible, the maintained posture throughout the fatiguing protocol was
not explicitly controlled for. Joint kinematic data may be limited in that an assumption was made
based on the selection of the drop landing used, that peak joint angles occurred while the feet
were in contact with the ground. Muscle activation data was limited as well, due to a small
sample size, consequence of poor goniometer data.
Summary
Abdominal fatigue may increase the risk of ACL injury by increasing associated
biomechanical factors. Explicitly, abdominal fatigue may increase the risk of ACL injury by
increasing hip flexion and increasing hip internal rotation. While inconsistent with literature as a
risk factor for ACL injury, abdominal fatigue significantly increased knee flexion (p=0.012 and
p=0.06 between pre-test and post-test1, and pre-test and post-test 2 respectively) but may be
indicative of subjects landing with a more quadriceps-dominant landing pattern. Additionally,
abdominal fatigue significantly increased quadriceps activation (p=0.015, pre-test to post-test 2).
The combination of kinematic and muscle activation results suggests that abdominal fatigue
causes subjects to land in a more quadriceps-dominant position, further straining the anterior
knee. These factors are correlated with ACL injury based on their contribution toward
controlling or increasing anterior tibial shear forces, and strain on the knee (Cowling & Steele,
2000; Kernozek et al., 2008; Ortiz et al., 2008; Shultz et al., 2009).
50
Chapter V
Summary, Conclusions, and Recommendations
Summary
This study observed the acute effects of abdominal fatigue on lower-extremity kinematics
and muscle activation patterns. Subjects were seated on the ground and were required to hold
their trunk at 40º relative to the ground, with their knees at 90º and their feet secured for as long
and as accurately as possible. Post-test drop-landings began 5 seconds after the completion of the
fatiguing protocol. There was a significant increase in peak knee flexion after abdominal fatigue.
However, non-significant increases in peak hip flexion and internal rotation were observed after
fatigue. Additionally, non-significant increases in peak knee valgus were accompanied by nonsignificant decreases in gluteus maximus EMG activity. Further, quadriceps activation
significantly increased after abdominal fatigue, with a corresponding non-significant increase in
quadriceps to hamstring activation ratio.
Conclusions
Abdominal fatigue may increase the risk of ACL injury in college aged females.
Specifically, holding the trunk at a 40° relative to the ground with the knees at 90° and feet
supported increased hip flexion, hip internal rotation and quadriceps activation, as well as
increases in knee valgus. Future research is necessary to understand how varying sources of
abdominal fatigue explicitly effect anterior shear joint forces at the knee.
51
Recommendations
Future Research. While this study identified abdominal fatigue as a risk factor for ACL
injury, further research is necessary to understand how abdominal fatigue alters anterior shear
joint forces at the knee. Anterior shear joint forces at the knee have been consistently referenced
as a key factor in ACL injury (Cowling & Steele, 2000; Kernozek et al., 2008; Ortiz et al., 2008;
Shultz et al., 2009), and as such, forces through the knee should be measured to better understand
how abdominal fatigue influences the forces causing ACL injury. Additionally, because this
study implemented a fatiguing protocol which likely recruited the hip flexors, further studies
should observe the effects of isolated abdominal fatigue protocols. In regards to muscle
activation patterns specific to ACL injury, EMG activity should be recorded for the other
quadriceps and hamstring muscles to provide a more complete activation profile of these muscles
after abdominal fatigue. Further research is necessary to identify the length of time abdominal
fatigue alters lower extremity kinematics and altered muscle activation patterns. Future research
should investigate if effects of abdominal fatigue differ between trained and untrained
individuals.
Practical Applications. Results from this study suggest that training to increase core
endurance may decrease the risk of ACL injury. For example, using abdominal exercises during
a warm-up for athletic events or physical activity, may be increasing the risk for injury under
some circumstances. Specifically, performing abdominal fatiguing exercises prior to plyometric
activities may be predisposing individuals to injury. Alternatively, learning and enforcing proper
landing mechanics to more efficiently absorb the anterior shear forces going through the knee
may help prevent ACL injuries in athletes and individuals. Specifically, learning to land with
greater hip extension and external rotation, as well as knee varus may reduce the risk of ACL
52
injury, in conjunction with increasing the strength and recruitment of the hamstrings to increase
the Q:H ratio.
53
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Appendices
Appendix A:
Informed Consent
Abdominal fatigue and Lower extremity kinematics
Purpose and Benefit:
Researchers of this study are interested in the interaction of abdominal fatigue and motions of the
hip and knee during a drop landing. Females are at an increased risk for anterior cruciate ligament (ACL)
injuries in the knee. As such, the results of this study will identify additional risk factors and potentially
recognize abnormal motions of the lower extremity that can help decrease the high risk of ACL injuries in
females.
I UNDERSTAND THAT:
1) This experiment will involve the completion of a series of tasks that include a brief, lowintensity warm-up on a cycle ergometer, two maximal abdominal contractions, multiple drop
landings, a sustained abdominal hold, and a brief low-intensity cool-down. Additionally, this
experiment will require me to be instrumented with electrodes on my abdomen and reflective
markers at various joints on your body. My participation will require approximately 60
minutes of my time.
2) There are minimal anticipated risks or discomfort associated with participants. Participants
may experience low grade abdominal soreness up to 48 hours post involvement. A possible
benefit of participation in this study may include understanding of the research process.
3) My participation is voluntary. I am able to withdraw from the study at any point in time
without penalty.
4) All information is confidential. My signed consent form will be kept in a locked cabinet
separate from any information tying me to this study. Only the primary researcher will have
access to any data gathered in this study. My name will not be associated with any
performances or relevant data to the study.
5) My signature on this form does not waive my legal rights of protection.
6) This experiment is conducted by Eryn Murphy under the supervision of Dr. Jun San Juan.
Any questions that you may have about the experiment or your participation may be directed
to the investigators at (360)650-2336
If you have any questions about your participation or your rights as a research participant, you
can contact the WWU Human Protections Administrator (HPA), (360)650-3220. If during or after
participation in this study you suffer from any adverse effects as a result of participation, please
notify the researcher directing the study or the WWU Human Protections Administrator.
I have read the above description and agree to participation in this study.
________________________________________________________________________
Participant’s Signature
Date
__________________________________________________
Participant’s PRINTED NAME
Note: Please sign both copies of the form and retain the copy marked “Participant”
66
Appendix B: Protocol Checklist
Subject ID: MT2015_
Date:
/
/
Informed consent signed
YES / NO
Informed consent understood, questions answered
YES / NO
Height
Weight
Gender
Does subject have a history of lower extremity or trunk injury?
Dominant Leg
R/L
YES / NO*
*if yes, subject must be excluded from the study
5 minute warm up on cycle ergometer completed
YES / NO
Discuss protocol:
1) We will demonstrate a drop landing and give you time to practice
2) Then we will get you fully instrumented with the EMG and joint markers
3) Once instrumented, we will perform the initial 3 drop landings
4) You will then be further instructed on the fatiguing protocol before performing the abdominal
hold
5) Immediately after the abdominal fatigue protocol, you will begin the final 3 drop landings.
Demonstration of a single drop landing completed
YES / NO
Subject offered time to practice drop landing
YES / NO
“Step off with your right foot, and step back onto the block (backward) with your left foot”
Instrumentation completed
YES / NO
EMG sensors on dominant side
□
□
□
□
□
□
External oblique
Superior Rectus abd.
Inferior Rectus
Glute Max
Vastus Medialis
Semitendinosus
MVCS performed
YES / NO
Qualisys markers placed
YES / NO
Explanation of fatigue protocol:
YES / NO
“For this fatigue protocol, we will have you lay against this wedge with your head against it, arms
across your chest. We will count down: 3-2-1-Go. On go we will pull this wedge out from behind you 2
inches, and lower it 20 degrees. Your goal is to maintain that initial posture for as long as possible,
until your trunk touches the wedge. As soon as your back touches the wedge, stand up and
immediately stand on the box for your final 3 drop landings. Once on top of the box, wait for the “Go”
command to begin your first landing, remembering to step off the block with your right foot, and back
up onto the block with your left foot. DO YOU HAVE ANY QUESTIONS?”
NOTES:
67
Appendix C: Raw data
Joint angle data for peak knee flexion and valgus, as well as hip flexion and internal rotation
(mean±SEM) at pretest, posttest-1 and posttest-2 measures.
KNEE FLEXION
KNEE VALGUS
HIP FLEXION
HIP INTERNAL
ROTATION
PRETEST
85.00 ± 2.71
10.66 ± 0.98
67.98 ± 3.98
POSTTEST-1
89.80 ± 2.60
10.99 ± 1.05
64.20 ± 4.09
POSTTEST-2
89.98 ± 2.54
11.04 ± 1.09
71.42 ± 3.89
5.65 ± 1.15
7.35 ± 1.07
7.68 ± 1.08
Average activation profiles for the rectus femoris (RF), biceps femoris (BF), gluteus maximus
(GM) and activation ratio between the RF and BF (Q:H ratio), (mean±SEM). Activations are
normalized to a maximal voluntary contractions.
Q:H RATIO
RF ACTIVATION
BF ACTIVATION
GM ACTIVATION
PRETEST
POSTTEST
3.07 ± 2.09
18.62 ± 1.90
14.07 ± 5.85
17.55 ± 3.50
4.03 ± 3.42
27.06 ± 3.76
9.88 ± 1.97
16.87 ± 3.78
Peak activation profiles for the rectus femoris (RF), biceps femoris (BF), gluteus maximus (GM)
and activation ratio between the RF and BF (Q:H ratio), (mean±SEM). Activations are
normalized to a maximal voluntary contractions.
PEAK
Q:H RATIO
RF ACTIVATION
BF ACTIVATION
GM ACTIVATION
PRETEST
3.86 ± 2.72
36.56 ± 3.44
26.55 ± 11.69
41.96 ± 11.41
68
POSTTEST
5.67 ± 5.74
55.39 ± 6.34
17.37 ± 4.06
34.19 ± 9.69