Influence of a Marching Snare Drum System on  
Joint Kinematics, Electromyography, and Contact Pressure 
 
by 
 
Andrea Michelle Sumner 
 
 
 
 
A dissertation submitted to the Graduate Faculty of 
Auburn University 
in partial fulfillment of the 
requirements for the Degree of 
Degree of Doctor of Philosophy 
 
Auburn, Alabama 
May 6, 2012 
 
 
 
 
Keywords: load carriage, EMG, contact pressure, 
         kinematics, kinetics 
 
 
Copyright 2012 by Andrea Michelle Sumner 
 
 
Approved by 
 
Wendi H. Weimar, Chair, Associate Professor of Kinesiology 
David D. Pascoe, Professor of Kinesiology 
Richard D. Good, Professor of Music 
Joellen M. Sefton, Assistant Professor of Kinesiology 
  
ii 
 
ABSTRACT 
 
 
 Each year thousands of students participate in marching band.  Many participate 
by carrying a flag, baton, woodwind or brass instrument, or drum.  The drumline is of 
particular interest due to the unique and restricting nature of the instrument: a load carried 
anterior to the body with the only points of contact on the body at the shoulders and 
across the abdominal area.  The influence of load carriage research on gait and joint 
mechanics, muscle activity, and contact pressure can be seen in the ever changing 
backpack designs and recommendations of load mass; however, there still remains a lack 
of research on marching band load carriage.  The purposes of this investigation were: 1) 
to determine the influence of a standard marching snare drum system (SDS) and a 
modified marching snare drum system (MDS) on trunk and lower extremity kinematics 
while standing and marching; 2) to determine the muscular demand placed on the erector 
spinae, of individuals wearing a SHS and MHS while standing and marching; and 3) to 
determine the contact pressures at the shoulders and abdominal region of individuals 
wearing the SDS and MDS systems. 
The results show that the SDS during stationary tasks decreased hip flexion, and 
increased trunk flexion, ankle plantarflexion, ES EMGAVG and EMGPEAK as a result of 
the changes in posture, and contact pressures at the shoulders.  When the participants 
wore the SDS while marching, the results showed increased trunk flexion, and increased 
ES EMGAVG and EMGPEAK as a result of the changes in trunk angle.  When implementing  
iii 
 
a lumbar belt there were similar significant findings at the joints and muscle activity level 
from the unloaded condition only, suggesting that the implementation of the lumbar belt 
did not alter the marching kinematics of the performer.  The MDS condition however, did 
result in lower contact pressure values at the shoulders and higher pressures at the 
abdominal region, indicating a redistribution of pressure and load to the belt, hips, 
abdomen, and pelvic girdle and off of the shoulders.  Therefore it can be concluded that 
implementing a lumbar belt to a SDS system ameliorates skin contact pressure and 
discomfort at the shoulders, potentially decreasing the occurrence of an injury, and 
anecdotally providing relief to the participants.   
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
iv 
 
ACKNOWLEDGMENTS 
 
 
 The author would like to thank Dr. Wendi Weimar for not only introducing her to 
the field of Biomechanics, but for the dedication to her academic and professional 
development and maturity.  Thanks to Dr. Weimar?s countless hours of guidance, 
wisdom, patience, and support, I was able to push myself academically to higher 
standards which allowed me to be the educator and researcher that I am today.   
 The author would like to thank the members of her committee, Dr. David Pascoe, 
Dr. Richard Good, and Dr. JoEllen Sefton for their time, patience, and dedication to my 
development throughout this process and my academic career at Auburn University.  The 
author would also like to thank Dr. David Shannon for his willingness to serve as the 
outside reader, and Dr. Sara Wolf for her assistance, advice, and support.  
  The author would also like to thank the members of the Sport Biomechanics 
Laboratory: Jay Patel, Braden Romer, John Fox, and Jared Rehm for the tireless help, 
hard work, and support throughout the data collection process.  To the members of the 
Auburn University Marching Band Drumline, the author is very thankful for their 
patience and willingness to participate. 
 Most importantly, the author would like to thank her parents, Gordon and Vicki, 
for their unwavering support and encouragement throughout her entire educational 
career.  Without their love, belief, and patience during the long journey, this would not 
have been possible.  I love you both TMAM.  
v 
 
Table of Contents 
 
Abstract  ........................................................................................................................................ ii 
Acknowledgments ....................................................................................................................... iv 
List of Tables ............................................................................................................................. viii 
List of Figures  ............................................................................................................................. ix 
List of Abbreviations  ................................................................................................................. xii 
Chapter I: Introduction .................................................................................................................. 1 
 Purpose ............................................................................................................................ 15 
 Hypotheses  ..................................................................................................................... 16 
 Limitations  ..................................................................................................................... 17 
 Delimitations  .................................................................................................................. 17 
 Definition of Terms......................................................................................................... 18 
Chapter II: Review of Literature  ................................................................................................ 20 
 Section 1: Biomechanical Effects of Posterior & Combined Anterior/Posterior           
                  Load Carriage................................................................................................. 21 
 Section 2: Biomechanical Effects of Anterior Load Carriage  ....................................... 27 
 Section 3: Muscle Activity & Contact Pressure ............................................................. 32 
 Section 4: Lumbar Belt Implementation  ........................................................................ 43 
 Section 5: Summary ........................................................................................................ 49 
Chapter III: Methods  .................................................................................................................. 53 
vi 
 
 Participants ...................................................................................................................... 54 
 Setting  ............................................................................................................................ 54 
 Materials ......................................................................................................................... 55 
 Instrumentation  .............................................................................................................. 61 
 Design & Procedures ...................................................................................................... 69 
 Experimental Design & Data Analysis  .......................................................................... 73 
Chapter IV: Results  .................................................................................................................... 75 
 Section 1: Participant Demographics  ............................................................................ .76 
 Section 2: Joint Kinematics ............................................................................................ 76 
 Section 3: Erector Spinae Muscle Activity ..................................................................... 85 
 Section 4: Contact Pressure ............................................................................................ 90 
Chapter V: Discussion   ............................................................................................................. .93 
 Section 1: SDS & MDS Conditions on Joint Kinematics ............................................... 94 
 Section 2: SDS & MDS Conditions on ES Muscle Activity .......................................... 99 
 Section 3: SDS & MDS Conditions on Contact Pressure ............................................. 102 
 Section 4: Summary & Future Research ....................................................................... 105 
References   .............................................................................................................................. .109 
Appendix A: Preliminary Medical Questionnaire (PMQ) ........................................................ 116 
Appendix B: Informed Consent ................................................................................................ 117 
Appendix C: ANOVA Table for the Effect of Load Carriage on Joint Kinematics ................. 120 
Appendix D: ANOVA Table for the Effect of Load Carriage on Joint Kinematics ................. 121 
Appendix E: ANOVA Table for the Effect of Load Carriage on ES sEMG ............................ 122 
Appendix F: ANOVA Table for the Effect of Load Carriage on Peak Contact Pressure ........ 123 
vii 
 
Appendix G: Contact Pressure at Both Shoulders during the Pre5 Measurement .................... 124 
Appendix H: Contact Pressure at Both Shoulders during the Pre55 Measurement .................. 125 
Appendix I: Contact Pressure at Both Shoulders during the Post55 Measurement .................. 126 
Appendix J: Contact Pressure at the Abdominal Region during the Pre5 Measurement ......... 127 
Appendix K: Contact Pressure at the Abdominal Region during the Pre55 Measurement ...... 128 
Appendix L: Contact Pressure at the Abdominal Region during the Post55 Measurement ..... 129 
Appendix M: Raw sEMG Signal of the ES for a Complete Gait Trail during the Unloaded  
            Condition............................................................................................................ 130 
 
Appendix N: Raw sEMG Signal of the ES for a Complete Gait Trial during the SDS  
           Condition............................................................................................................. 131 
 
Appendix O: Raw sEMG Signal of the ES for a Complete Gait Trial during the MDS  
           Condition............................................................................................................. 132
viii 
 
List of Tables 
 
 
Table 1: Kinematic Challenges of the SDS  ............................................................................... 51 
Table 2: Names and Positions of Markers used in the Plug-in-Gait Model  .............................. 61 
Table 3: Participant Demographics  ............................................................................................ 76 
Table 4: Effect of Loading Conditions on Joint Kinematics ...................................................... 94 
Table 5: Effect of Loading Conditions on sEMG of the ES ..................................................... 100 
Table 6: Effect of Loading Conditions on sEMG  .................................................................... 103 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
ix 
 
List of Figures 
 
 
Figure 1: Snare Drum (Pearl, 2010)  ........................................................................................... 19 
Figure 2: Snare Drum Carrier (Randall May, 2010)  .................................................................. 19 
Figure 3: Length of the Carrier and Height of the Drum  ........................................................... 55 
Figure 4: Distance between the Bottom of the Bracket Arm and Top of Drum  ........................ 56 
Figure 5: Snare Drum Carrier Shoulder Frame-Sagittal View ................................................... 56 
Figure 6: Snare Drum Carrier Shoulder Frame-Posterior View  ................................................ 56 
Figure 7: Snare Drum Carrier-Abdominal Plate-Posterior View  .............................................. 57 
Figure 8a: Lumbar Belt Construction-Dimensions  .................................................................... 59 
Figure 8b: Lumbar Belt Construction-Dimensions .................................................................... 59 
Figure 9: Lumbar Belt Design-Belt Corner Angle  .................................................................... 59 
Figure 10a: Lumbar Belt Design-Neoprene Layers & Foam Placement  ................................... 60 
Figure 10b: Lumbar Belt Design-Neoprene Layers & Foam Placement  ................................... 60 
Figure 11: Lumbar Belt Design-Tubing Used for Framework  .................................................. 60 
Figure 12: Lumbar Belt Design-Velcro Placement & Attachment  ............................................ 60 
Figure 13: Anterior View of the Plug-in-Gait Marker Locations  .............................................. 62 
Figure 14: Posterior View of the Plug-in-Gait Marker Locations  ............................................. 62 
Figure 15: Sagittal View of the Plug-in-Gait Marker Locations on Leg  ................................... 62 
Figure 16: Frontal View of the Plug-in-Gait Marker Locations on Lower Leg  ........................ 62 
Figure 17: Trunk Angle (Attwells et al., 2006)  ......................................................................... 63 
x 
 
Figure 18: Hip Angle (Attwells et al., 2006)  ............................................................................. 63 
Figure 19: Knee Angle (Attwells et al., 2006)  ........................................................................... 63 
Figure 20: Ankle Angle (Attwells et al., 2006)  ......................................................................... 63 
Figure 21: Vicon Motion Capture Camera Laboratory Setup  ................................................... 64 
Figure 22: Functional Phases of Gait (Kirtley, 2008)  ................................................................ 65 
Figure 23: Noraxon MyoSystem? Tethered Surface EMG Receiver ......................................... 66 
Figure 24a: Noraxon Telemyo? 2400R ?World Wide Telemetry Receiver ............................... 66 
Figure 24b: Noraxon? Telemyo 2400T V2 Wireless Transmitter .............................................. 66 
Figure 25: sEMG Electrode Placement for the Erector Spinae  ................................................. 66 
Figure 26: Tekscan I-Scan System? (v 6.10.1) ........................................................................... 68 
Figure 27: Tekscan I-Scan System? Pressure Sensor #9801 ...................................................... 68 
Figure 28: Tekscan I-Scan System? Pressure Sensor #9830  ..................................................... 68 
Figure 29: Pressure Sensor Locations on the Carrier ................................................................. 71 
Figure 30: Pressure Sensor Data Collection ............................................................................... 71 
Figure 31: Effects of Loading Conditions on TAAVG during Stationary Trials .......................... 78 
Figure 32: Effects of Loading Conditions on TAAVG during Dynamic Trials ............................ 79 
Figure 33: Effects of Loading Conditions on HAAVG during Stationary Trials .......................... 80 
Figure 34: Effects of Loading Conditions on HAAVG during Dynamic Trials  .......................... 81 
Figure 35: Effects of Loading Conditions on AAAVG during Stationary Trials .......................... 82 
Figure 36: Effects of Loading Conditions on AAAVG during Dynamic Trials  .......................... 83 
Figure 37: Effects of Loading Conditions on KAAFF  ................................................................. 84 
Figure 38: Effects of Loading Conditions on AAAFF  ................................................................. 85 
Figure 39: Effects of Loading Conditions on ES EMGAVG Before a One Minute Exercise  ...... 87 
xi 
 
Figure 40: Effects of Loading Conditions on ES EMGAVG After a One Minute Exercise ......... 88 
Figure 41: Effects of Loading Conditions on ES EMGPEAK During Stationary Trials  .............. 89 
Figure 42: Effects of Loading Conditions on ES EMGPEAK During Dynamic Trials................. 90 
Figure 43: Effects of Loading Conditions on SHLDRPEAK & ABSPEAK Contact Pressure ........ 92 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
xii 
 
List of Abbreviations 
 
 
AAAVG  Average Ankle Angle 
ABSPEAK  Peak contact pressure at the abdominal/pelvis area 
BOS  Base of Support 
COG  Center of Gravity    
COGperson  Center of Gravity of an individual 
COGpack  Center of Gravity of the load carriage device 
COGsys  Center of Gravity of the person + pack 
ES  Erector Spinae 
HAAVG  Average Hip Angle 
KAAVG  Average Knee Angle 
LSPEAK  Left Shoulder Peak contact pressure 
MDS  Modified Marching Snare Drum System-Drum + Harness + Lumbar Belt 
RSPEAK  Right Shoulder Peak contact pressure 
SHLDRPEAK Peak contact pressure at the shoulders 
sEMG  Surface electromyography 
SDS  Standard Marching Snare Drum System-Drum + Harness 
TAAVG  Average Trunk Angle 
 
 
1 
 
CHAPTER I 
INTRODUCTION 
 
 Music is a universal language that can evoke every emotion, regardless if it 
resonates from a church choir, a national orchestra, or a college marching band.   In order 
to achieve the desired performance, musicians must train.  Train as a musician to perform 
the music and, in the case of marching bands, train as an athlete to perform the on-field 
performance and while carrying an instrument.   
Marching music presentations require not only musicianship but also demanding 
movements, postures, and instrument weight carriage.  This can provide a challenge for 
some of the larger awkward instruments or those whose carriage position alters the 
location of the center of gravity of the body.  It is the load carriage demands of marching 
music presentations that this study will investigate.  Specifically, the biomechanical 
effects associated with wearing a snare drum carrier and drum, and how the 
implementation of a lumbar belt to the snare drum carrier influences these effects. 
There are many load carriage systems available on the market today such as 
backpacks, carrying bags, rucksacks, holsters, and belt clips.  Some of these systems 
require that the load be carried on only one shoulder, supported by both shoulders, or 
supported on the pelvis and hips.  The present study will investigate the dual shoulder 
method and implementing a device to place a portion of the load on the pelvis, and as 
such will not address the issues associated with asymmetrical loading.   
 2 
Load Carriage 
Load carriage allows an individual to transport an additional mass whether it is on 
the anterior part, posterior part, or sides of the body, or through the use of a carrying 
device (i.e., backpack, rucksack, or carrier).  Examples of load carriage that are found in 
the research include: hiking backpacks (Cook & Neumann, 1987; Goh, Thambyah, & 
Bose, 1998), school bookbags (Devroey, Jonkers, Becker, Lenaerts, & Spaepen, 2007), 
military load bearing equipment (Ling, Houston, Tsai, Chui, & Kirk, 2004; Martin & 
Nelson, 1986), and ergonomic/lifting (Anderson, Meador, McClure, Makrozahopoulos, 
Brooks, & Mirka, 2007; Cham & Redfern, 2004).  However, even with the abundance of 
individuals participating in marching bands throughout the country, there is little research 
on the influence of load carriage in marching bands.  Nearly absent are investigations into 
the anterior load of a standard marching snare drum system (SDS-snare drum and 
carrier).  The few studies available on marching bands have only examined whether or 
not participation in that activity can be a substitute for a physical education class (Strand 
& Sommer, 2005) or examined a survey on marching band injuries sustained during a 
season (Mehler, Brink, Eickmeier, Hesse, & McGuire, 1996).  The characteristics of the 
SDS are unique and include: the carrier has minimal adjustment capabilities; the mass of 
the snare drum will always be located anteriorly to the individual; there is a lack of a hip 
or back support on the carrier; and there are only three points of contact throughout the 
entire carrier (upper trapezius region, upper back-scapular region, and an abdominal 
plate).  Research on backpacks, rucksacks, and ergonomic tasks provide the foundation 
for understanding the biomechanical influence of load carriage systems on the human 
 3 
body, and for that reason, understanding the impact of the SDS and beginning to establish 
this line of research will be inferred from research on posterior and anterior load carriage.     
 
Biomechanical Effects of Load Carriage Mass  
Posterior & Combined Anterior/Posterior Load Carriage 
One of the most popular methods of load carriage in the current era is the 
bookbag.  It is estimated that over 40 million individuals, primarily students, use 
bookbags as a means to carry personal belongings, work-related materials, and 
particularly school supplies and books.  It is recommended that the maximum about of 
load carriage placed on the body should not exceed 20% of body mass (BM) (Malhotra & 
Sen Gupta, 1965; Pascoe & Pascoe, 1999).  The student population however, can end up 
carrying loads weighing more than the 20% of the individuals? BM from the textbooks 
alone, before additional supplies, lunches, and other miscellaneous materials are 
calculated into the total load. It is these large loads that are believed to contribute to the 
significant kinematic and kinetic changes that have been associated with back pain or 
injury for those who utilize this method of load carriage (Devroey et al., 2007; Hong, Li, 
& Fong, 2008; Malhotra & Sen Gupta, 1965; Pascoe & Pascoe, 1999).   
Another popular method of load carriage, and one that is as specific to a particular 
activity is a hiking pack.  Hiking packs are unique from bookbags by utilizing either an 
internal or external frame and often use lumbar belts which provide the pack with an 
additional stabilizing and load absorption component, (Reid, Stevenson, & Whiteside, 
2004; Smith, Ashton, Bohl, Clark, Metheny, & Klassen, 2006).  The literature provides 
evidence that the addition of support structures (lumbar belt and/or frames) aid in the 
 4 
transfer of forces from the body to the frames (Reid et al., 2004), shift the center of 
gravity of the system (COGsys) location of the load closer to the body (Bloom & 
Woodhull-McNeal, 1987; Goh et al., 1998), transfer forces from one part of the body to 
another more suited for load carriage (i.e., from the shoulders to the pelvis and hips), and 
decrease the muscular demand to support the load (Holewijn, 1990; Harman, Han, 
Frykman, Johnson, Russell, & Rosenstein, 1992).   
Military rucksacks are examples of a posterior load carriage system, as are hiking 
packs and school backpacks and therefore cause individuals to demonstrate similar 
responses.  Military rucksacks utilize rigid external frames, similar to hiking packs, 
which allow for a more stable load carriage system for the soldiers (Bloom & Woodhull-
McNeal, 1987).  The difference between the load carriage systems is that a soldier does 
not have the option of managing the items carried and therefore does not have control 
over the weight of the pack.  While the external frames are light weight, they still 
contribute to the overall mass of system in addition to the mission-required equipment.  A 
military rucksack has a mass of approximately 21 kg empty; therefore with the addition 
of the required supplies the mass of the system can reach well over the recommended 10-
20% BM that should be placed on the body (BAE Systems, 2010; Devroey et al., 2007; 
Hong, Li, & Fong, 2008; Malhotra & Sen Gupta, 1965; Pascoe & Pascoe, 1999).  
Military load carriage can also consist of loads being placed on the anterior and posterior 
aspects of the body, including the utility vest worn under the rucksack , body armor, and 
double packs (Birrell & Haslam, 2009; Harman et al., 1994; Kinoshita, 1985).  Research 
assessing the effect of the double packs (half the mass anterior, half the mass posterior) 
revealed similar kinematic results to anterior and posterior load carriage methods, 
 5 
specifically; decreases in stride length and an increase in stride frequency (Knapik, 
Reynolds, & Harman, 2004). 
An individual?s body alignment adapts to increased loads as a means of 
alleviating the impact of the load.  The trunk adopts a more forward leaning position to 
accommodate the posterior shift of the individual?s center of gravity (COGperson); a 
forward neck position is adopted to assist the increased trunk flexion in shifting the 
COGsys forward, there is decreased balance control, and therefore may lead to subsequent 
lower back pain (Al-Khabbaz, Shimada, & Hasegawa, 2008; Devroey et al., 2007; Goh et 
al., 1998; Pascoe, Pascoe, Wang, Shim, & Kim, 1997).  Furthermore, gait analyses reveal 
that increasing the mass of the backpack results in decreased stride length and single 
support time, and increased stride frequency and double support time (Kinoshita, 1985; 
Wang, Pascoe, & Weimar, 2001).  It should be noted that there the center of gravity can 
be defined as the COG of the load/pack (COGpack), the COG of the person (COGperson), 
and the COG of the pack/person system (COGsys).       
The static and dynamic responses to posterior and combined anterior/posterior 
load carriage include a posterior and vertical shift in COGsys, the trunk adopts a more 
forward lean, there is increased double-support time, muscular activity, vertical and 
horizontal force distribution on the body, and subsequent loading of the lower extremities 
(Birrell and Haslam, 2009, 2010; Goh et al., 1998; Grimmer, Dansie, Milanese, Pirunsan, 
& Trott, 2002; Ling et al., 2004).  Although carrying loads on the back has been the 
primary load carriage method throughout history (Knapik, Harman & Reynolds, 1996), 
there are other methods that allow an individual to transport items, for example: loads 
 6 
placed only on the anterior portion of the body, identical to the demands seen in marching 
drumlines.   
 
Anterior Load Carriage 
When carrying an external (i.e., a rifle, front packs, or office crates) or internal 
(i.e., pregnancy) load anteriorly, the static posture goals remain the same: perform the 
required task while limiting the amount of COGsys excursion from that of an unloaded 
location, limiting excess muscular demand, and decreasing forces exerted on the lower 
extremities and within the vertebral column (Legg & Mahanty, 1985).  In addition to 
consistency in static posture goals, many of the same outcomes noted during the gait 
cycle in posterior load carriage hold true for anteriorly positioned loads.  These findings 
include a more abrupt loading phase during gait, a shift in the COGsys vertically and 
anteriorly, and an increase in hip and knee flexion angle (Anderson et al., 2007; Cham & 
Redfern, 2004).   
Pregnancy research supports these findings and has shown that the anterior shift 
of the COGperson in addition to the increase in body mass contributes to an increase in 
anterior pelvic tilt and double support time, and decreases in hip extension and single 
support time (Foti, Davids, & Bagley, 2000).    The result of those kinematic adjustments 
require an increase in muscular demand, and may provide an explanation of the 
musculoskeletal/overuse issues seen in anterior load carriage.  Myung and Smith (1997) 
investigated anterior load carriage and indicated that stride length was influenced by the 
mass of the load; however, the load investigated was 40% BM, well above previous 
research testing loads.  Although the research to suggest a relationship between anterior 
 7 
load mass and stride length were studied at much lower values (20% BM), the overall 
findings in this research are indicative of the real-world and therefore should be taken 
into account. 
 
Biomechanical Effects of Vertical Load Carriage Location  
In addition to the load carriage mass, research on load carriage includes 
investigating the result of the load carriage?s vertical position.  It is recommended that the 
load be positioned as close to the COGperson, in both the vertical and horizontal directions.  
Increasing the distance of the load away from the COGperson in any direction can 
compromise an individual?s balance (Qu & Nussbaum, 2009) and increase strain on the 
body.  Carrying a load just 30 cm horizontally from the hip, for example, has been shown 
to cause a trunk flexion moment of up to 20 Nm.  The result is an equal but opposite 
extension moment to counteract the forward tilt, and increased muscular demand (Cham 
& Redfern, 2004). 
Higher load placement during static conditions may allow an individual to shift 
the COGsys more easily in an attempt to maintain balance; however, in order to maintain 
that balanced state, there is an increase in the muscular demand due to the larger moment 
that is created.  Maintaining a higher COGsys also means a decrease in stability (less 
allowable movement within the base of support (BOS)), which actually helps initiate 
movement and is therefore preferred for level walking.  Lower placement for static 
conditions is thought to cause more adjustments by the body to maintain balance, yet a 
lower COGsys is more stable and can result in less muscular demand due to a smaller 
 8 
moment arm.  A lower COGsys resulting in better balance during movement has led other 
investigators to believe it is better for walking on uneven terrain. 
Southard and Mirka (2007) researched dynamic conditions and recommend 
lumbar height placements during level walking conditions for posterior loads to maintain 
balance and decrease muscular demand.  Bobet and Norman (1984) agree and 
recommended placing the load in the lower thoracic region during uneven terrain 
movements as high placements can result in increased moments and muscle activity at 
the lower back.  Knapik et al. (2004) concur that lower load placement is better suited for 
uneven terrain by lowering the COGsys closer to the ground, and higher load placements 
are better suited for level walking as it helps initiate the movement.  The conclusion, 
vertical load placement recommendations are complicated and unclear.  
Load placement height issues differ between anterior/posterior loads, 
static/dynamic movements, even/uneven terrain, and research in this area is less abundant 
that the literature on load carriage mass.  Regardless, the method of carrying (anterior 
versus posterior) load mass (% body mass) and location (thoracic versus lumbar regions) 
are all factors affecting gait, stability measures, and muscular adaptations.   
 
Muscle Activity & Contact Pressure 
 Load carriage will predominately affect the primary point(s) of contact, whether 
the load is placed on the legs or arms (Knapik et al., 2004), or for the purposes of this 
study, the shoulders.  In addition to affecting the kinematics of the individual during 
stationary and dynamic tasks, consideration must be given to the outcomes of muscle 
activity and skin contact pressures due to load carriage. 
 9 
Surface Electromyography 
Any deviation from the normal, unloaded, anatomically neutral posture results in 
a change in muscular demand (Bobet & Norman, 1984; Hamilton, Weimar, & Luttgens, 
2008).  Deviations that occur during load carriage situations can be from increased load 
mass, changes in the amount and location of force being exerted on the body, changes in 
load placement, or a combination of any of these factors.  Researchers not only found that 
increasing load mass causes an increase in muscular activity, but also supported the 
recommendation that loads not exceed 15% of body mass (Al-Khabbaz et al., 2008; 
Devroey et al., 2007).  In addition to investigating the changes in muscle activity due to 
load carriage mass and location, researchers also considered the length of time that the 
load was being carried.  The results indicated that prolonged load carriage can lead to the 
excess muscular tension that can potentially lead to muscular pain and fatigue (Bobet & 
Norman, 1984; Hong et al., 2008).    
While gait and balance demands associated with load carriage involve the whole 
body, this discussion will be limited to the muscles identified in the research as being 
most involved with posterior and anterior load carriage: i.e., erector spinae (ES), rectus 
abdominis (RA), and upper trapezius (UT) (Al-Khabbaz et al., 2008; Bobet & Norman, 
1984; Devroey et al., 2007; Hong et al., 2008; Schiffman, Bensel, Hasselquist, 
Gregorczyk, & Piscitelle, 2006).  Bobet and Norman (1984) investigated the difference 
between the trapezius (TRAP) and ES for loads carried at the ear lobe level (first cervical 
vertebrae) and xiphoid process (tenth thoracic vertebrae).  These researchers found the 
midback load placement resulted in less muscle activity in both the ES and TRAP 
muscles, with the erector spinae muscles exhibiting less activity overall than the TRAP 
 10 
(Bobet & Norman, 1984).  Al-Khabbaz et al. (2008) found that the RA activity, but not 
the ES, changed significantly at loads of 10% BM or greater in backpacks.  This occurred 
due to the posterior load on the trunk causing an increase in trunk extension.  Although 
the ES are trunk stabilizers, the significant muscle activity occurred in the trunk flexors in 
order to counteract the extension moment.  Hong et al. (2008) found that backpacks 
weighing 15-20% BM significantly increased the muscular demand of the lower TRAP 
due to the increased demand on the shoulders from the backpacks. 
Load carriage has also been found to increase trunk stiffness (increased level of 
muscular contraction).  Muscular stiffness can occur as a result of the joint(s) being 
hindered in movement (Birrell & Haslam, 2008) or due to a decrease in stability as a 
result of an external load (Shape, Holt, Saltzman, & Wagenaar, 2008).  During load 
carriage situations, a decrease in stability of the trunk can occur due to the additional 
flexion, extension, lateral, and rotational moments that occur as a result of the load 
carriage system shifting the COGperson away from the unloaded position. The result is an 
increase demand placed on the musculature that can lead to increased pain, early onset of 
fatigue, and injuries.   
Load mass and placement has been shown to influence muscle activity.  Increased 
muscle activity occurs due to the body being loaded as well as the additional adjustments 
that are made as the body attempts to compensate for the extra load.  For example, the 
responses occur in the musculature opposite to the joint action (i.e., increased abdominal 
activity due to the trunk flexion as a response to spine extension caused by the load).  
Research to date has centered on the points of the body with which the load makes 
contact (i.e., the trunk), instead of the moments that are created as a compensatory 
 11 
mechanism.  The research on the muscle activity responses of the lower extremities to 
load carriage is limited, but the trends found by Al-Khabbaz et al. (2008) indicate that the 
lower extremity musculature will show increased activity in an assisted role.  Therefore, 
more research is needed to determine whether load carriage influences the lower 
extremity muscles (i.e., gastrocnemius and soleus), or whether the trunk musculature is 
solely responsible for managing the increased demands. 
 
Contact Pressure 
In addition to investigating the kinematics and muscle activity associated with 
load carriage, skin contact pressure has also been examined as it has been considered a 
limiting factor during load carriage (Holewijn, 1990).   Increased load mass impacts not 
only the kinematic responses of the body but also causes an increase in skin contact 
pressure (Holewijn, 2000; Mackie, Stevenson, Reid, & Legg, 2005) and can have 
deleterious consequences on the individual.  Skin contact pressure of more than 105mm 
Hg (14kPa) results in irritation and redness; and if exposed for more than two hours, may 
lead to subcutaneous edemas (Husain, 1953).  High pressures are also indicative of the 
impact the load has on the structures below the surface of the skin, for example the 
superficial nerves of the brachial plexus.  If these structures are pinched, damaged, or 
compressed, the result is a condition known as Rucksack Palsy.  This pressure decreases 
the efficacy of the neuromuscular signal from reaching the target muscle resulting in 
weakness and/or loss of function (Goodson, 1981; Knapik et al., 1996, 2004).   
In an effort to combat the effects of load carriage not only on the joints, muscle 
activity and contact pressure, researchers have investigated the implementation of a 
 12 
lumbar belt.  In a study particularly related to skin contact pressure, the researcher noted 
a 36% increase in pressure as the load mass increased within the pack system without a 
lumbar belt (Holewijn, 1990).  This study was able to show a relationship between load 
mass and load carriage type, and the significance of a lumbar belt.   
 
Lumbar belt Implementation 
In addition to suggesting lighter masses and alternate loading positions, there is 
also research investigating the impact of implementing supportive devices, such as a 
lumbar belt.  LaFiandra and Harman (2004) investigated the forces exerted on the upper 
and lower back and found that the addition of a lumbar belt helped redistribute the 
location of the forces on the body to the pelvis.  Redistributing these forces to the pelvis 
is beneficial because the hips are three times less sensitive to pressure, are located closer 
to the COGperson position, and are the preferred (anecdotally) placement for load carriage 
(Knapik et al., 2004; Scribano, 1970).  Sharpe et al. (2008), expanding on the work by 
LaFiandra, Wagenaar, Holt, and Obusek (2003), examined the effect of implementing a 
lumbar belt on the coordination between the pelvis and trunk movements and found that 
the belt improved the stability of the pelvic-thorax coordination pattern by decreasing the 
muscular demand of the trunk stabilizers.  Essentially, the lumbar belt allowed for less 
restrictive pelvic-trunk rotation, while providing stability between the segments and 
thereby decreasing the muscular demand.    
Implementing a lumbar belt has been shown to be beneficial in the relationship 
between muscle activity and skin contact pressures.  Holewijn (1990) for example found 
that 10kg carried in a frameless pack resulted in a peak skin pressure of 203mmHg at the 
 13 
shoulders; that same mass carried in a pack with a flexible frame and lumbar belt resulted 
in decreased TRAP muscle activity and a peak pressure of only 15mmHg.  While 
Holewijn did not ascribe the decrease in muscular activity to the lumbar belt alone, a 
subsequent study showed that without the use of a lumbar belt, there was a 44-47% 
increase in shoulder pressure (Mackie et al., 2005).  
Loads produce tension on the anterior and superior aspect of backpack shoulder 
straps and thereby contribute to the increased pressure exerted on the shoulders.  The 
amount of tension and pressure is influenced by load mass, load location, tightness of 
straps, and use or non-use of a lumbar belt.  Without the use of lumbar belt, both the skin 
contact pressure and increased muscle activity become significant and potentially task 
limiting factors.  With the aid of a lumbar belt a portion of the load can be redistributed to 
the waist, decreasing the pressure and muscular demands of the shoulders and providing 
for a more efficient method of load carriage.   
 
Summary 
Load carriage has many ramifications on the body such as: decreased balance and 
stability (shifting the COGperson more to the edge of the BOS), increased muscular 
demand, and increased potential for injury.  In particular, research has demonstrated that 
load carriage decreases stride length (Birrell & Haslam, 2009; LaFiandra et al., 2003), 
single support time (Attwells, Birrell, Hooper, & Mansfield, 2007; Wang et al., 2001), 
hip extension angle (Anderson et al., 2007; Cham & Redfern, 2004), and increases 
double support time (Goh et al., 1998; Kinoshita, 1985; Wang et al., 2001), stride 
frequency (LaFiandra et al., 2003; Martin & Nelson, 1986), and knee flexion angle 
 14 
(Anderson et al., 2007; Devroey et al., 2007; Holt et al, 2003; Kinoshita, 1985).  Overall, 
these adaptations were seen during bookbag, military, hiking, and ergonomic research, all 
areas where there are fewer performance constraints on the individual.  Although there 
are some similarities between military and marching band load carriage (i.e. regimented 
step lengths and cadences), and yet there is a dearth of research within the marching 
arena.  Marching music performers are also constrained by the body posture, step 
frequency, and having a fixed load.  These limitations highlight the inability to 
automatically apply the solutions found in load carriage research to marching load 
carriage tasks. 
The findings of the noted research are expected to be confirmed in the present 
study if the marching snare drummer were not to be constrained in a number of ways.  
Specifically: (a) there is no horizontal adjustment on the SDS; (b) The vertical location of 
the drum is set for the proper playing position first and the aesthetics of the drum line 
second.  The height of the individual has little influence on the height of the drum; (c) 
There is a standardized step size, no matter the leg length of the individual (and often a 
standardized stride frequency); (d) The rigidity of the carrier results in higher contact 
areas as the shape of the carrier cannot be adjusted, and (e) the position of the drum 
(directly anterior to the thighs limiting hip and knee angles) prevents this population from 
employing the protective and stability mechanisms seen in the more common load 
carriage situations.   
 
 
 
 15 
Statement of the Problem 
 Currently there are thousands of individuals participating in marching bands, yet 
there is little research on the topic.   Within the literature on marching bands, the most 
frequently discussed topic is whether or not marching band can be considered a 
replacement for physical education classes (Cowen, 2006; Erdmann, Graham, Radlo, & 
Knepler, 2003; Strand & Sommer, 2005).  The medical issues encountered throughout a 
marching season have also been investigated (Jones, 1983; Mehler et al., 1996) but still 
remains a topic less frequently investigated.  In order to advance the research in the area 
of marching band load carriage, the first step will be to determine if the findings of the 
proposed study are concurrent with previous literature on load carriage.  The next step 
will be to investigate the influence of the addition of a hip support to the SDS, and the 
role of the lumbar belt in alleviating the negative side effects anecdotally seen in 
marching band load carriage. 
 
Statement of Purpose 
 The purposes of this research are: 1) to determine the influence of a SDS on trunk 
and lower extremity kinematics while standing and marching; 2) to determine the contact 
pressures and muscular demand placed on the erector spinae, of individuals wearing a 
SDS; and 3) to investigate the influence of a lumbar belt (MDS-modified SDS ) on joint 
kinematics, muscular activity, and contact pressures.  A standard snare drum system 
(SDS) consists of the initial carrier (Figure 1) and drum (Figure 2), and an MDS consists 
of the SDS with the addition of a lumbar belt. 
 
 16 
Hypotheses 
Primary Objective-To determine the postural demands of a marching snare drum system 
1) Evaluate the joint kinematics during a load-carrying task using both an SDS and 
MDS in marching drumline individuals. 
H01: During the stationary movement condition, the SDS will result in increased 
trunk extension (smaller trunk flexion angle), hip flexion, knee flexion and 
plantarflexion.  
H02: During the dynamic movement conditions, the SDS will result in increased 
trunk extension, decreased hip flexion, increased knee flexion and increased 
dorsiflexion. 
 
Secondary Objective-To determine the muscle activity and contact pressures associated 
with wearing a marching snare drum system  
1) Evaluate the muscle activity of the erector spinae during a load-carrying task 
using both an SDS and MDS in marching drumline individuals. 
H01: The participants will demonstrate higher levels of muscle activity while 
wearing the SDS, during both the stationary and dynamic movement 
conditions. 
 
2) Evaluate the pressures at the shoulders and abdominal/pelvic region of marching 
band individuals during a load-carrying task using both the SDS and MDS during 
the stationary movement condition.  
H01: The participants will demonstrate higher contact pressures at both the 
shoulder and abdominal region from the SDS. 
 17 
H02: The participants will demonstrate reduced contact pressures at the 
shoulders and increased contact pressures at the abdominal region from 
the MDS condition. 
 
Limitations 
The limitations for the present study include: 
1. Self-reported health status was used. 
2. The current marching snare drum system used by Auburn University Marching 
Band Snare Line consisted of a Pearl? Championship Maple FFX-1412 Snare 
Drum (Figure 1) with a Randall May? Aluminum Tube Snare Carrier (Figure 2).  
3. The data collection took place at the beginning of the participants? season, when 
the drums are being worn on a regular basis. 
 
Delimitations 
The delimitations for the present study include: 
1. Participants were required to wear retroreflective markers bilaterally and surface 
electrodes unilaterally. 
2. Participants were between the ages of 19-25, members of the Auburn University 
Marching Band Drumline, and have at least one year (high school or college) 
experience performing with a marching snare or tenor drum. 
3. The participants were assessed in a laboratory versus on the marching field or 
road (parade comparison). 
 18 
4. The drum height was placed at an Auburn University standard, determined by the 
director, to be 10? below the horizontal for each individual.  
 
Definition of Terms 
Anatomical Joint Angle: The angle between two adjoining segments. 
 
Base of Support: The area in which a part of the body is in contact with the supporting 
surface within the outline of both feet (Holbein & Chaffin, 1997). 
 
Center of Gravity (COG): The point of a body about which all mass is said to act.  The 
point of a body about which the system is balanced. 
 
Kinematics: A branch of mechanics that investigates pure motion, without regard to the 
forces that produce that motion.  Kinematic variables include position, velocity and 
acceleration.   
 
Moment Arm: The perpendicular distance from the line of force to the point about which 
moments will be summed. 
 
Plantarflexors: Muscles of the lower limbs documented to maintain upright stance as 
well as being involved in the propulsion phase of the gait cycle.  Muscles included are the 
soleus, gastrocnemius, tibialis posterior, flexor digitorum longus, flexor hallucis longer, 
peroneus brevis and tertius. 
 19 
Postural Stability: The body?s capability to control the center of gravity within specific 
boundaries of space (limits of stability), enabling the individual to maintain in upright 
posture (Shumway-Cook & Woollacott, 1995).  
 
Standard Marching Snare Drum System (SDS): The load carriage (LC) apparatus used 
for this study and by the Auburn University Marching Band Drumline include the Pearl? 
(Pearl? Corporation, Nashville, TN, USA) Championship Maple FFX-1412 Snare Drum 
(7.67kg) with a Randall May? (Randall May? International, Inc., Irvine, CA, USA) J-
Rodless Contour Hinge? Monoposto? Snare Carrier (0.91kg) (Figures 1 & 2 
respectively).  
     
Figures 1& 2. Snare Drum (Pearl, 2010) and Snare Drum Carrier (Randall May, 2010). 
 
Surface Electromyography (sEMG): A technique for evaluating and recording the timing 
and magnitude of voluntary skeletal muscle activity.  EMG is performed using an 
instrument called an electromyograph that produces an electromyogram (record of data) 
(MacIntosh et al., 2006; Winter, 1990).  
 
 20 
CHAPTER II 
REVIEW OF LITERATURE 
 
Thousands of marching band musicians engage in demanding and strenuous 
movements regularly while practicing and performing, and a select group of musicians 
perform these movements while carrying large cumbersome loads.   However, the 
influence of a marching snare drum system (SDS) on the biomechanics of a marching 
band member has received little attention in the literature.  Therefore, this load carriage 
demand will be investigated in order to provide information about this population and the 
load carriage techniques being utilized.  Specifically, this study will investigate how 
wearing a snare drum carrier and drum effects the trunk and lower extremities; and the 
influence of a lumbar belt on the biomechanical demands placed upon the bodies of 
marching band drummers.   
 This chapter is divided into five sections.  Section one will summarize the 
previous work in posterior and combined anterior/posterior load carriage.  Section two 
will describe the static and dynamic effects of anterior load carriage.  Section three will 
describe the effects of load carriage on muscle activity and skin contact pressure.  Section 
four will describe the impact of implementing a lumbar belt to a load carriage system.  
Finally, section five will summarize the pertinent findings of the previous literature to the 
present study. 
 21 
Section 1: Biomechanical Effects of Posterior & Combined Anterior/Posterior  
Load Carriage 
During unloaded conditions, the line of gravity of the head, arms, and trunk 
(HAT) is located anterior to the lumbosacral (LS) joint creating a more extension 
dominated posture (Bobet & Norman, 1984).  This places an initial increased demand on 
the extensor musculature in order to maintain an erect posture and maintain the COGsys 
within the base of support (BOS).  It should be noted that there is the center of gravity of 
the load/pack (COGpack), the COG of the person (COGperson), and the COG of the 
pack/person system (COGsys).  During posteriorly loaded conditions these adaptations are 
even more pronounced and include: increased muscle activity (Al-Khabbaz et al., 2008; 
Devroey et al., 2007), increased motion of the trunk and lower extremities (Chow, Leung, 
& Holmes, 2007; Devroey et al., 2007; Holt et al., 2003; Seven, Akalan, & Yucesoy, 
2008), and a shift in the center of gravity (COGsys) of the combined load and body system 
(Cook & Neumann, 1987; Holt et al., 2003).  Likewise, posteriorly loaded gait trials also 
produce movement adaptations that include:  decreases in single support time (SST) 
along with increases in double support time (DST) (Kinoshita, 1985), decreased stride 
lengths and increased stride frequencies (LaFiandra et al., 2003), changes in force 
distribution (Bloom & Woodhull, 1987; Reid et al., 2004), and subjective reports of pain 
and discomfort (Moore, White, & Moore, 2007).   
 During static posterior loaded conditions, although less movement is required to 
move the COGsys within the BOS, a higher COGsys is also less stable because of the 
longer moment arm in the sagittal plane, particularly during dynamic movements.  
However, depending on the proximity relative to the axis of rotation, a lower COGsys 
 22 
tends to require additional compensation by the body (i.e., movement at the trunk, hips, 
and ankles) (Bloom & Woodhull, 1987; Devroey et al., 2007).  It is by understanding 
these principles of static load carriage and the possible outcomes from significant 
changes in posture that a foundation for understanding the effects of load carriage to 
dynamic situations can be provided.   
 Posterior loading during dynamic conditions alter the COGsys excursion amounts 
and knee kinematics (Holt, Wagenaar, LaFiandra, Kubo, & Obusek, 2003), influences the 
forces measured at the shoulders and waist (LaFiandra & Harman, 2004; Mackie et al., 
2005), and can also result in significant changes in joint kinematics (Harman. Frykman, 
Knapik, & Han, 1994; Kinoshita, 1985; Smith et al., 2006).  Both Holt et al. (2003) and 
LaFiandra et al., (2003) discovered that during unloaded conditions when walking speed 
is increased, the body adopted a longer stride length, increased stride frequency, 
increased knee flexion, and increased transverse pelvis and trunk rotation in order to 
maintain the required speed.   When the body was loaded however, walking and 
maintaining speeds greater than 1.0 m/s resulted in decreases in pelvic and thoracic 
rotation, decreased stride length, and an increase in stride frequency.  The lack of 
movement between the pelvis and trunk is primarily due to increases in trunk and lower 
extremity stiffness, possibly from muscle co-activation in an effort to stabilize the pelvis 
and minimize the rotational forces on the already challenged upper body (Birrell & 
Haslam, 2009; Sharpe et al., 2008).   
Similar to LaFiandra et al., (2003), Birrell and Haslam (2009) found that as the 
load mass increased, particular within a framed pack, the pelvic girdle rotation decreased.  
The decreased pelvic girdle rotation may have occurred not only as a protective 
 23 
mechanism described above, but also due to the restricted arm swing associated with 
carrying a rifle during one of the test conditions, (Birrell & Haslam, 2009).  Natural arm 
movements modulate the excursions of the COGsys (Birrell & Haslam 2008), therefore it 
can be assumed that limiting this movement, as seen during military load carriage with a 
rifle or a marching band snare drummer, can impede this response and cause increased 
demands elsewhere in the body and result in limited joint range of motion (ROM).    
Military load carriage is also unique in that it places a demand on all sides of the body for 
load carriage (anterior and posterior).  This occurs in the load carriage vests worn under 
the rucksacks, the webbing system vests worn, canteen belts, and body armor, and is 
referred to as combination load carriage.  One attempt to combat the decreased stability 
of the pelvis due to load carriage, while still allowing the joints to move in an efficient 
manner, is to implement a lumbar belt.  This topic will be discussed further, but as was 
noted by Sharpe et al. (2008) that loading participants with a lumbar belt equipped 
backpack resulted in significantly more pelvic girdle-to-trunk rotation pattern in addition 
to a more stable pattern in the participants? gait.   
Dynamic studies on combination load carriage have revealed similar kinematic 
results to separate anterior and posterior load carriage methods such as a decrease in 
stride length (Birrell & Haslam, 2009; Harman et al., 1994), decreased trunk flexion 
(Harman et al., 1994) and an increase in stride frequency predominately due to the body 
being externally loaded.  Kinoshita (1985) found that as the load mass increased, 
regardless of a posterior load or combination load design, the period of DST increased, 
knee flexion increased, and SST decreased.  The significant change in knee flexion aided 
in the shock absorption during the loading phase, and correlated with the increased DST 
 24 
which may facilitate in the transition of weight between the feet.  Kinoshita (1985) found 
that as the load mass increased in the backpack the participants utilized increased step 
widths which is indicative of decreased pelvic girdle rotation and a decrease in balance.  
The doublepack system resulted in less forward trunk lean and less altered gait 
kinematics.  This difference between the double pack and single backpack have some 
researchers (Birrell & Haslam, 2010; Kinoshita, 1985; Knapik et al., 1997; Motmans, 
Tomlow, & Vissers, 2006) considering full body load distribution rather than one sided 
loading as the best option.   
Both males and females will be included in the present study; therefore a brief 
review of the gender differences during load carriage tasks is appropriate.  A majority of 
the load carriage research uses males only as participants.  Presumably because of the 
predominance of men in the military, and because most load carriage devices are built 
based on male anthropometry.  However, according to the U.S. Census (2010), there has 
been an increase in the number of females in the military: 197,900 active duty in 2008 to 
213,823 in 2010 (Department of Veterans Affairs, 2010; infoplease?, 2011).  Females 
have structural differences that affect the ability to carry loads in the same manner as 
men.  For example, females typically have a shorter stride length and must increase stride 
frequency during ruck marches to maintain the same speed as the longer striding males.  
Females also generally have less muscle mass, less bone density, and increased joint 
laxity during the menstrual cycle; all of which may contribute to muscular or skeletal 
injuries during a mission (Ling et al., 2004).  Females also have a wider pelvis that is not 
well suited to the narrow lumbar belts found in most rucksacks or hiking packs.  This 
mismatch can result in discomfort at the pelvis as well as excessive movement of the 
 25 
pack (Birrell & Haslam, 2009; Ling et al., 2004).  Another difference between males and 
females that should be taken into consideration is the general smaller size of the female 
and thus the increased load mass relative to body mass that must be carried.   
In response to these differences, Ling et al. (2004), and Martin and Nelson (1986) 
chose to investigate the kinematic and temporal responses of females to different load 
carriage situations.  Ling et al. (2004) investigated the influence of a military load 
carriage system (vest and rucksack) of 0, 20, 30, 40, and 50 pounds on seven young-adult 
females.  Each condition consisted of three, pre-march walking trials at a speed of 4.827 
km/hr across a 40 foot walkway; and a 60 minute simulated march trial on a treadmill at 
4.827 km/hr; followed by another three post-march walking trials across the walkway.  
Ling et al. (2004) noted that the female participants responded to the increased load 
similar to men, with increased DST, decreased SST, increased hip and trunk flexion, and 
increased vertical COGsys movement.  The females also displayed decreased knee flexion, 
thus increasing the stiffness of the lower leg at impact and provided less shock 
absorption.  Although there were no significant changes in stride length or stride 
frequency, the result did indicate similar trends shorter stride lengths and faster stride 
frequencies seen in previous studies.  The insignificance could have occurred as a 
consequence of the researcher?s stringent significance level (? < 0.005).  It should be 
noted that one participant was unable to complete all the trials due to the substantial 
discomfort over the pelvic girdle from the pack.  The issue was due to an anatomical 
difference of having a small waist and wide pelvis that could not be fitted comfortably 
with the pack.  In addition, three participants had to make modifications (i.e., custom 
pads added to the points of contact on the pelvis) in order to sufficiently shift the load 
 26 
from the shoulders to the hips for better comfort.  These results provide further evidence 
of the need to design a load carriage system that is more adaptable across the sexes and 
body sizes.      
Martin and Nelson (1986) compared eleven male and eleven female ROTC 
students while walking over level ground at 1.78 m/s with the same absolute mass 
conditions (9, 17, 29, and 36 kg).  The results revealed that both males and females had 
decreased stride lengths and increased stride frequencies as the load increased.  
Moreover, the females had overall shorter stride lengths and increased stride frequencies 
than the males for all conditions.  Both groups had decreased swing times as the load 
increased, yet the females had significantly shorter swing times from the males for the 
heaviest loads (17, 29, and 36 kg).  The females showed an increase in DST as the load 
mass increased, with significant changes from the males occurring for the two heaviest 
loads.  The males demonstrated less consistent changes in the DST during the loading 
conditions, but did yield a significant difference for the 9 kg and 36 kg loads.   Both the 
males and females had significant trunk flexion for the 29 and 36 kg loads, with the 
females maintaining larger trunk angles overall, suggesting no effect of sex.  Overall, the 
females had less trunk lean than the males and significantly less lean during the first three 
conditions.  It was hypothesized that less trunk lean may have occurred as a result of 
already having a lower COGperson, and therefore the females were less influenced by the 
increased rucksack load. 
Research has indicated that placing an external load posteriorly on the body or on 
both the anterior and posterior side of the body result in similar coping strategies.  
Although the posterior and combined literature provide the foundation for studying 
 27 
marching drum load carriage, it also reveals the current gap in the marching band load 
carriage literature.  For example, the constraints of the marching gait limit adaptations 
such as decreases in stride length (Harman et al., 1994), decreased single stance time 
(SST) (Hong & Li, 2005), decreased pelvic rotation (Devroey et al., 2007), increases in 
stride frequency (Harman et al., 1994), and increased double stance time (DST) (Hong & 
Li, 2005).  However, as unique as marching band load carriage tasks are, the literature on 
primarily anterior load carriage provides the structure for the changes in joint kinematics 
at the trunk, hip, knee, and ankle (Goh et al., 1998; Harman et al., 1994; Shasmin, Abu-
Osman, Razali, Usman, & Wan Abas, 2005), changes in knee angle to aid in shock 
absorption (Kinoshita, 1985), and increased muscle activity (Devroey et al., 2007) that 
are anticipated to occur during this study. 
 
Section 2: Biomechanical Effects of Anterior Load Carriage 
The kinematic and kinetic results of carrying a load anteriorly are similar to 
posterior and combined anterior/posterior load carriage.  Specifically, the following 
alterations have been identified: a shift in the COGsys vertically and anteriorly (instead of 
posteriorly), an increase in trunk extension (instead of flexion), and increases in hip and 
knee flexion angle (Cham & Redfern, 2004; Anderson et al., 2007).  Although there is 
limited research regarding anterior load carriage, there is research in the field of 
ergonomics that can provide understanding of the effects of being loaded or carrying 
loads on the anterior portion of the body.  Anderson et al. (2007) for example, 
investigated anterior load carriage on posture while eleven healthy college-aged males 
stood still for 6 seconds and walked at a self-selected pace (ranging from 1.9-3.3 mph) on 
 28 
a level treadmill for 15 seconds.  The participants performed trials with two loads: a 
hand-held barbell, mass equaling 20% of participants? maximum voluntary isometric 
elbow flexion, and a 14kg bucket.  Both objects were held in the front on the body at 
heights in line with the knuckles, elbows, and shoulders.  Overall, the trunk ROM while 
walking was 1.5? more than standing at the ninth (T9) and twelfth thoracic (T12) and 
third lumbar vertebrae (L3).  Holding the barbell at all three heights resulted in increased 
trunk extension at T9 and T12 vertebrae during both the standing and walking conditions.  
When standing with the 14kg bucket at shoulder height, trunk extension not only 
occurred for both T9 and T12 vertebrae, but the L3 vertebrae as well.  Similar to the 
barbell condition, carrying the bucket at shoulder height resulted in the largest trunk 
extension angle during gait, with similar ROM differences.  The results of this study 
support previous literature on posterior load carriage that suggests that the trunk will 
adopt a leaning posture when loaded.  This lean will move the COGsys closer to the center 
of the BOS in order to try and maintain a balanced, upright posture.  During posterior 
load carriage an anterior lean is observed; in response to an anterior load, a more 
posterior lean is utilized.  Recall that an unloaded body presents with an extended posture 
due to the natural anterior carrying position of the head and trunk, maintained by the 
isometric contractions of the tonic musculature.  Therefore, an anteriorly loaded body 
will require the trunk to extend even farther, putting a greater demand on this 
musculature. 
Both Cham and Redfern (2004) and Myung and Smith (1997) investigated the 
impact of anterior load carriage and the surface of the floor on gait kinematics.  Myung 
and Smith (1997) specifically investigated stride length as well as heel velocity due to a 
 29 
correlation between certain heel velocities and increased potential to slips/falls.  Ten 
college-aged males, walked at a self-selected pace around a 25 m circular walkway under 
five loading conditions (unloaded, 20% BM in a container, 40% BM in a container, 18 kg 
container ?held against the body?, and 24 kg container ?held against the body?) over four 
floor surfaces (oily vinyl tile, dry stainless steel, oily plywood, and dry ceramic tile).  
Heel contact during normal gait ranges from 10-20 cm/s and can increase up to 50.8 cm/s 
during a slip.  The researchers discovered that the heel velocity for dry floors was slower 
than the oily floors regardless of the loads, and the stride lengths were longer for the dry 
floors indicating the participants had better balance on the dry floors.  Slippery surfaces 
required individuals to adopt a shorter stride length in order to maintain as much contact 
with the floor as possible and to keep the feet as close to the COGperson as possible.  These 
adaptations resulted in increased stride frequency during fixed speeds, which in turn 
related to a higher heel velocity.  These findings confirmed previous research that larger 
loads shorten the stride length, but the predominate influence on the gait was the floor 
surface.  At a fixed speed the decreased stride length allowed for quicker transfer of load 
to the adjacent supporting limb, thus, decreasing time in single support and resulting in an 
increase in balance.   
Cham and Redfern (2004) recruited ten healthy males to walk up a ramp 
positioned at 0, 5, and 10?, during three loading conditions (no load, 2.3 kg, and 6.8 kg).  
The load conditions consisted of weights placed in an office crate held bilaterally at waist 
level (elbows flexed approximately 90?) at a self-determined horizontal distance from the 
body.  The results indicated significant increases in plantarflexion, knee flexion, hip 
flexion, trunk extension, as well as decreases in stance duration, and heel velocity.  The 
 30 
increased plantarflexion, hip and knee flexion were affected by the changes in ramp angle 
across all angles, but only between the 0 and 6.8 kg loading conditions.  While the 
changes in stance duration decreased linearly with the increased load mass, these 
outcomes occurred to a greater extent during the 0? angle compared to the 10? ramp.  
This occurs because the incline causes the foot to make contact sooner, therefore 
contributing to the increased knee and hip angles in order to swing the leg through 
quicker than when on level ground.  Cham and Redfern?s (2004) study suggests that the 
changes in these kinematic variables may help an individual decrease the likelihood of 
slipping.  The significance of which is because of the various surfaces that a marching 
band individual will come across throughout the season including hills on campus and 
wet parade routes.  However, once the slip occurs chances of falling may increase due to 
the increased inertia of the person/load and by having the arms occupied with the load. 
Research has shown similar to posterior load carriage, pregnant individuals have 
increased anterior pelvic girdle tilt and double support time, and a decrease in hip 
extension and single support time (Foti et al., 2000).  Foti et al. (2000) tested fifteen 
healthy women during the third trimester of pregnancy and one year post partum.  The 
participants were asked to perform repeated self-selected paced walks covering 
approximately 12 m on level ground.  The kinematic changes found by Foti (2000) 
during the pregnant trials were believed to be compensations as the pregnant subjects 
attempted to maintain as non-pregnant a gait as possible.  For example, pelvic width 
(width between the anterior superior iliac spines) and ankle separation increased during 
DST.  Compensation included increases in the hip adduction moment in order to keep the 
feet more centered under the body thus avoiding wide BOS indicative of a waddling gait.  
 31 
A larger ankle plantarflexion moment was used as a means of increasing the propulsive 
forces in order to maintain a forward/linear gait and overcome the increases in body 
mass.   
Anterior loads cause an individual?s balance and stability to be compromised as 
indicated by the changes in kinematic and kinetic responses (Abu Osman & Ghazali, 
2002; Anderson et al., 2007; Cham & Redfern, 2004; Foti et al., 2000).  One method of 
compensating for these changes is to increase flexion at the hip and knee, resulting in the 
ability to swing the leg forward to place the foot on the ground quicker.  In doing so, the 
result is a longer stance phase, longer periods of DST, and a better transition of the load 
to the opposite limb (Kinoshita, 1985).  Holbein and Redfern (1997) tested fifteen males 
while unloaded and carrying an 11.4 kg box during a functional limits of stability test.  
The researchers noted that increased flexion at the hip and knee lowers the COGsys 
thereby increasing stability and extending an individual?s stability limits.  When carrying 
a load a set distance horizontally from the pelvis a flexion moment at the trunk is created 
(Anderson et al., 2007).  Causing an equal and opposite extension moment at the trunk by 
the erector spinae (ES) and hip extensors at the pelvis.  These moments influence the 
muscle activity seen in anterior load carriage, placing a higher demand on the 
musculature of the hip, pelvis, and trunk simultaneously in order to contribute to the 
increased demand for stability (Anderson et al., 2007).   
Lee and Lee (2003) investigated twelve male participants while unloaded and 
carrying a 12 kg load above the head, at shoulder height, and at knuckle height with a 
wide foot placement (45 cm apart) and a narrow foot placement (20 cm apart).  Each 
participant stood on a platform, barefoot, in an erect posture for 7 seconds followed by a 
 32 
functional limits of stability test.   This study indicated that the wider foot placement 
lowered the COGsys possibly contributing to increased stability. 
Load carriage requires postural adaptation to maintain balance and translate the 
load to a new location.  Research to date has focused mainly on posterior load carriage 
and anterior movements.  However, marching band activity requires the participants to 
translate in all directions.  In particular, snare drummers are often required to perform a 
?crab-walk? which is a cross-over step gait that permits the performer to move laterally 
while still facing the director and continuing to direct the sound forward.  This project 
will address the unique conditions associated with marching band snare drummer load 
carriage including anterior load carriage, restricted arm movements and translation in 
multi-directions. 
 
Section 3: Muscle Activity & Contact Pressure  
The majority of load carriage literature has focused on the kinematic changes in 
body posture and gait dynamics as stand-alone responses.  However, changes in joint and 
gait kinematics influence how the muscles must respond to the loading demands on the 
body.  The following section will present the relevant literature associated with 
electromyography and contact pressures during load carriage. 
 
Surface Electromyography 
Posterior loads causes the COGperson to shift backwards creating an extension 
moment in the trunk (Cook & Neumann, 1987).  The body responds by flexing the trunk 
to move the COGsys closer to its unloaded position, resulting in increased levels of the 
 33 
trunk flexors (RA).  Al-Khabbaz et al. (2008) utilized a frameless bookbag during a static 
task  to determine the level of RA and ES activity when nineteen college-aged males 
stood for fifteen seconds during an unloaded and 10, 15, and 20% BM loaded condition.  
The researchers found an increase in only the rectus abdominus (RA) activity and no 
significant change in the ES activity, even though the participants in this study showed a 
backward inclination in posture as the load mass increased.  The trunk posture inclined 
backwards to the level of 3.37, 3.02, and 3.90? for the 10, 15, and 20% BM load 
conditions respectively, resulting in increases in the right RA to levels of 35, 105, and 
157% and 20, 65, and 110% for the left RA.  These findings show that posteriorly 
loading the trunk results in trunk extension and that the most significant changes in RA 
activity occurred at 20% BM, supporting previous research that loads of 20% BM should 
be avoided.  Although the ES activity levels were not significant, possibly due to the role 
of the ES as a trunk stabilizer and being utilized daily, the ES trends showed increasing 
activity as the load increased.  The ES and RA were active simultaneously and showed 
increased levels of activity as the load increased.  Circumstances in which the body will 
utilize the current antagonist musculature simultaneously in support of the agonist 
muscles of that particular task is known as coactivation (Latash, 2008).   
Coactivation occurs to stiffen the joint in an attempt to compensate for the effects 
of a perturbation, and as a safety mechanism of the central nervous system (Latash, 
2008).  In the case of being externally loaded, coactivation will occur in order to help to 
maintain the stability of an individual.  It is important to note that even though 
coactivation is considered a protective mechanism, the increased levels of muscle activity 
may contribute to the onset of muscular fatigue (Holewijn, 1990).    
 34 
In a study looking at the effect of load carriage and muscle activity, Bobet and 
Norman (1984) investigated the muscle activity of the ES and upper trapezius (UT) of 
eleven males while unloaded and while carrying a fixed mass of 19.5 kg in a framed 
backpack.  The load was placed on the posterior portion of the trunk at ear level and at 
the height of the xiphoid process, while each person walked 5.6 km/h along a level 90 m 
long course.  A change in muscle activity can occur as a result of a change in the posture 
of the body, as well as placing the body under an external load (Bobet & Norman, 1984).  
The results of this study indicated that loading the body (22-34% BM of participants) at 
heights corresponding to ear level and the xiphoid process increased the muscle activity 
in both the ES and UT.  Overall, the ES activity levels were lower than the UT, and 
between the two load placements, the lower placement (xiphoid process) produced the 
lowest sEMG values for each muscle group.  Positioning the load at ear level resulted in 
the body becoming less stable compared to the unloaded condition, and this instability 
was compensated for by the musculature.   
Motmans et al. (2006) investigated the effect of not only unframed backpacks, but 
shoulder bags (messenger bags), front packs, and double packs on the muscle activity 
levels of the two primary trunk stabilizers: RA and ES, bilaterally.  Nine male and ten 
female college students stood erect with the feet 15 cm apart for 30 seconds, while 
unloaded as well as carrying a shoulder bag, backpack, front pack, double pack, all 
loaded with 15% BM.  Similar to previous kinematic research, the double pack did not 
produce a significant difference from the unloaded position.  Furthermore, the findings of 
Motmans et al. (2006), supported the work of Birrell and Haslam (2010) that double 
packs promote a more upright stance similar to being unloaded, due to similarities in 
 35 
muscle activities of the RA and ES for the double packs and unloaded conditions.  During 
the backpack condition, the muscle activity of the ES decreased (77%) and RA increased 
(54%) while standing, as a result of the extension moment created by the load.  When 
loaded with the front pack, the participant?s response was to use the ES muscle group to 
move the COGsys in a position more centered within the BOS (Bloom & Woodhull-
McNeal, 1987; Bobet & Norman, 1984; Devroey et al., 2007).  This resulted in activity 
levels of more than twice the reference (unloaded) values: 206%.  The shoulder bag also 
showed a significant increase in ES activity, though not as high as the backpack or 
frontpack and also produced a significant difference between the contralateral (left) and 
ipsilateral (right) sides.  Overall the muscles on the contralateral side had the highest 
activity levels (Motmans et al., 2006), as a result of the participants laterally flexing the 
trunk opposite from the load as noted in previous literature (Pascoe et al., 1997).  
Devroey et al. (2007) also investigated the effects of wearing a frameless 
backpack while twelve male and eight female college students stood still for one minute 
and walked for five minutes loaded with 5, 10, and 15% BM.  The researchers found that 
while standing still, increasing the load mass to 10% and 15% BM caused increased trunk 
flexion, pelvic anteversion, and hip flexion.  While walking, increasing the load mass to 
15% BM resulted in a significant increase in trunk flexion, although there were also 
trends towards increased pelvic anteversion and hip flexion.  Devroey et al. (2007) 
hypothesized that the forward lean in this study occurred because the posterior location of 
the load caused the COGsys to shift backward, resulting in the kinematic compensations 
of the body to move the COGsys more within the BOS.  These joint kinematic changes 
also resulted in changes in muscle activity.  Activity of the RA and external obliques 
 36 
(EO) increased at 15% BM compared to 0% and 5%, while the ES muscles decreased at 
10% and 15% BM compared to 0% BM.  The decreased ES activity demonstrated at the 
lowest load that the coactivation of both the abdominal and back muscles as a protective 
mechanism was not occurring, but instead the load was being carried ?passively? 
according to the researchers (Devroey et al., 2007).  It is also possible that the decrease in 
the ES activity may have occurred as a result of 1) reciprocal inhibition or 2) increasing 
the forward lean of the body to allow the load to be distributed over more surface area 
(i.e., the pelvis and trunk), thus decreasing the demand/need for ES to be activated.   
Cook and Neumann (1987) also investigated the sEMG of the erector spinae 
during the stance phase of gait during an unloaded and loaded condition.  The researchers 
asked twelve young-adult males and twelve females to walk 30 m, at 1.3 m/s.  The load 
(10 and 20% BM) was placed in a wooden box, inside a framed backpack, carried 
contralaterally or ipsilaterally to the ES, and anteriorly and posteriorly to the chest.  The 
ipsilateral position yielded a decrease in the muscular demand with no change between 
the load mass conditions.  Both load conditions carried contralaterally to the electrodes 
on the ES resulted in a significant increase (11.69% MVIC) compared to the no-load 
condition, with a significant difference also noted between the load mass conditions 
(9.1% vs. 14.3% MVIC).  Of the four load conditions, the anterior position elicited the 
highest sEMG activity (11.93% MVIC) from the no-load condition as well as between 
the load mass conditions (9.9% vs. 14% MVIC, 10% and 20% respectively).  Overall, the 
female participants showed greater sEMG activity of the ES muscles during the loading 
conditions, with a 4.3% difference during the anterior load carriage position than the 
males (14.1% vs. 9.8% MVIC).  One possible explanation for the significant increase of 
 37 
muscle activity during the anterior loading condition was that the anterior condition 
required the use of both hands.  Restricting arm use was shown by Birrell and Haslam 
(2008) to prevent the horizontal motion of the trunk, resulting in greater excursions of the 
COGsys and a decrease in stability.  In order to compensate for the instability, the 
musculature must contract and is more likely to result in an earlier onset of fatigue or 
?fatigue-like effects? (Cook & Neumann, 1987). 
Hong, Li, and Fong (2008) not only looked at load mass on muscle activity, but 
also load duration in children carrying bookbags up to 20% BM.  The amount of time an 
individual wears a load is as influential as the load mass, load placement, and points of 
contact with the body.  Children and adults do not typically wear a pack of some kind or 
lift a load in industry for mere seconds; therefore it becomes advantageous to investigate 
the effect of load duration on an individual.  Fifteen male elementary-aged children were 
asked to wear a frameless school bookbag, unloaded (0% BM) or loaded with school 
supplies equaling 10, 15, and 20% BM, while walking on a level treadmill for 20 minutes 
at 1.1 m/s.  Surface EMG signals were collected throughout the duration of the walk for 
the UT, RA, and lower trapezius (LT).  As duration increased from zero to twenty 
minutes, there was a tendency of the muscle activity to increase.  A significant increase 
was found for the LT after 20 minutes for 0% BM, 15 minutes for 15% BM, and only 5 
minutes for 20% BM (fatigue levels were found at 15 minutes of the 20% BM).  These 
results provide further support that loads weighing 20% BM should be avoided (Devroey 
et al., 2007; Kinoshita, 1985) and duration should not last longer than 15 minutes (Hong 
et al., 2008).  
 38 
Harman et al. (1992) and Han et al. (1992) worked together on similar studies 
investigating the muscle activity of externally loaded individuals while altering either the 
load mass (Harman et al., 1992) or walking speed (Han et al., 1992).  Both groups of 
researchers asked fifteen physically-fit men to walk across a level platform while 
carrying loads of varying mass in a military style backpack.  Harman and colleagues 
(1992) utilized four different load conditions that equaled masses of 6, 20, 33, and 47 kg.  
The participants were asked to maintain a gait velocity of 1.3 m/s.  As a result of altering 
the amount of load being carried, Harman et al. (1992) found that average sEMG of the 
gastrocnemius (GAS) and quadriceps (Quad) increased as the load increased, and the 
average sEMG of the ES doubled at the 47 kg load.  These results may have occurred 
during the loading phase where the talocrural and tibiofemoral joints become 
eccentrically flexed, therefore requiring increased activity of the GAS and Quad 
respectively.  The increased activity of the ES was expected based on the load being 
placed on the trunk (Harman et al., 1992), but as a trunk stabilizer, may have only 
responded significantly for the excessive load mass.   
Han and colleagues (1992) utilized one load condition of 20 kg, but altered gait 
velocities to include 1.1, 1.3, and 1.5 m/s.  The researchers found significant increases in 
muscle activity of the GAS and Quad following a 36% increase in gait velocity (1.1 to 
1.5 m/s), as well as increases in the hamstrings (HAM) and tibialis anterior (TA) activity.  
However, there was no significant increase in the ES muscle activity regardless of the 
velocity, most likely due to the lower extremity musculature being primarily responsible 
of maintaining gait velocity (Han et al., 1992).  In addition, due to there being no change 
in load masses between the walking speeds, the ES simply did not need to change its 
 39 
activity level to handle the 20 kg load.  Both studies took the opportunity to collect 
additional gait kinematic and kinetic data, and found supporting results to the previous 
dynamic load carriage studies, for example: increased stride frequency, increased braking 
impulse, increased propulsive and medial forces, and increased % in DST. 
In addition to investigating the demands of carrying a load anteriorly to the body 
at various heights on trunk ROM and posture, Anderson et al. (2007) investigated the 
activity levels of the RA, EO, biceps brachii (BB), anterior deltoid (AD), and the ES at 
three vertebral levels (T9, T12, and L3) during seven load carrying tasks.  Eleven healthy 
college-aged males, were exposed to 2 load conditions (a hand-held barbell equaling 20% 
of participants? maximum voluntary isometric elbow flexion and a 14kg bucket) and 3 
height conditions (knuckle, elbow, and shoulder height).  Activity levels of the RA, EO, 
BB, AD and ES were recorded while holding the anterior loads, during a 6 second 
standing test and a 15 second walking test, at a self-selected pace on a level treadmill.  
Kinematic results demonstrated that while holding the barbell, trunk extension at the T9 
and T12 vertebrae were greatest while standing; however, the average trunk ROM at all 
three landmarks was 1.5? greater while walking.  In addition, carrying the barbell at 
shoulder height resulted in a larger trunk extension posture at the T9 and T12 vertebrae.  
The muscle activity during the barbell condition provided supporting evidence of the 
increased trunk ROM while walking.  There was also an increase in muscle activity of the 
ES at the T9 level while carrying the barbell at shoulder height.   In addition, the 
researchers observed that the normalized RA, EO, BB, and AD activity levels while 
walking were lower than all three ES levels (Anderson et al., 2007), indicating that the 
ES were primarily responsible for extending the trunk in order to shift the COGsys 
 40 
posteriorly to a more balanced location.  The other musculature were occupied with 
holding the load (BB and AD), while the RA and EO were contributing in a supportive 
role (i.e., coactivation).   
When standing with the 14kg bucket at shoulder height, trunk extension not only 
occurred for both the T9 and T12 vertebrae, but at the L3 vertebrae as well.  Similar to 
the barbell condition, carrying the bucket at shoulder height resulted in the largest trunk 
extension angle while walking, and an increase in the average trunk ROM.  As a result, 
the bucket condition also elicited a higher amount of muscle activity while walking, 
particularly for the ES at the T9 (35%), T12 (36%), and L3 (42%) levels, and the RA 
(132%).  The largest RA activity occurred while standing and holding the bucket at 
shoulder height, and while walking and holding the bucket at elbow height. 
The results of this study support the previous literature on posterior load carriage 
that suggests that the trunk will adopt a leaning posture when loaded in order to move the 
COGsys closer to the center of the BOS (Harman et al., 2004; Kinoshita, 1985; Smith et 
al., 2006).  Placing that COGsys closer to the center of the BOS will therefore allow an 
individual to maintain a more balanced and upright posture.  During load carriage 
situations, a decrease in stability of the trunk can occur due to the additional flexion, 
extension, lateral, and rotational moments that occur as a result of the load carriage 
system shifting the COGperson away from the unloaded position. The result is an increase 
demand placed upon the musculature that can lead to increased strain, onset of fatigue, 
and injuries.   
 
 
 41 
Contact Pressure 
Skin contact pressure can be considered one of the primary limiting factors in the 
ability to perform a load carriage task if the tension from the pack cannot be reduced 
(Holewijn, 1990).  Sangeorzan (1989) discovered that applying a pressure of 75mm Hg 
(10kPa) on skin directly over a muscle can reduce the circulation of blood.  A skin 
contact pressure of more than 105mm Hg (14kPa) was found to result in irritation, pale or 
dull skin, congestion of vessels, and if exposed for two hours, can induce subcutaneous 
edemas (swelling caused by fluid collection) and possible muscle fiber degeneration 
(Husain, 1953).  Doan (1998) collected both objective and subjective data on nine 
military load carriage systems from various countries and found subjective responses that 
support the claims that pressures above 14 kPa can be a limiting factor for the user.  
Twenty-eight soldiers volunteered to perform a variety of activities that tested balance, 
mobility, and agility, at four different activity stations for two of the nine load carriage 
systems.  At each station, the soldiers were asked to rate the load carriage system based 
on factors including the sense of the load being controlled, comfort, and amount of ROM 
of the arms and torso.  The soldiers completed the four station circuit for a total distance 
marched of 6 km.  Results showed that 95% of the soldiers reported discomfort when the 
average shoulder pressure exceeded 20 kPa, and 90% reported discomfort at 18 kPa.  It 
should be noted that the load that was being carried was 24 kg, approximately 30% BM 
of the average mass of the participants (Doan, 1998), which is well above the 
recommended 10-15% BM by previous research (Malhotra and Sen Gupta, 1965; Pascoe 
and Pascoe, 1999; Devroey et al., 2007; Hong et al., 2008).  The concern here is that the 
soldiers did not comment or notice any issues with the load carriage systems at the lower 
 42 
and still damaging levels of 10-14 kPa, indicating that some internal damage could be 
occurring before the individual is aware and protective mechanisms such as moving the 
pack to prevent blood occlusion can occur.  In a study investigating the impact of load 
mass and carrying mode, Holewijn (1990), demonstrated that skin contact pressure (kPa) 
was a physical parameter that could not only limit load carrying ability but load carrying 
duration as well.   
Husain (1953) performed a study of the legs of rats using a pressure cuff to 
determine the durations of pressure that could be withstood by an animal and the 
response it created.  It was determined that low pressures (100 mm Hg vs. 800 mm Hg) 
maintained for long periods of time can cause more damage than high pressures for short 
periods of time.  In addition, it was found that if the initial leg had pressure applied to it 
again three days after release, a much smaller pressure (50 mm Hg) applied for 1 hour 
caused muscular degeneration to begin to occur.  Based on these findings, Husain (1953) 
recommended that a substantial resting period should be utilized between loading 
conditions.  Husain (1953) also investigated 10 case studies of human patients with bed 
sores in order to document the duration in bed and the etiology of the musculature either 
beneath or near the bed sore.  The researcher determined that bed sores are associated 
with prolonged pressure (as early as one hour), nearness to a bony prominence, or a 
decreased ability of the patients to withstand normal pressure thresholds as a result of 
multiple medical conditions.  Fortunately, Husain (1953) also noticed that if it were 
possible to remove the excess pressure (i.e., repositioning the patient in bed) a 
hyperaemic effect resulted, where the blood flow was restored.  Non bed-ridden 
individuals do this every day as second nature.  For example, sitting during a two-hour 
 43 
lecture, a student will change position within the chair to alleviate the uncomfortable 
sensation in the buttocks.  
 The inability to relieve excess pressures has detrimental effects, the extent of 
which depends on the location (i.e., bed sores at the superficial level, decreased blood 
flow at the vessel level).  If the pressure occurs at the nervous system level, the result can 
be decreased sensations and loss of function.  Goodson (1981) determined that high skin 
contact pressures from tight fitting straps, such as those often used in an attempt to place 
the backpack load closer to the trunk, can affect the superficial nerves in the brachial 
plexus.  This is known as Rucksack Palsy and can result in symptoms anywhere along the 
arm depending on the site of the pressure.  If these nerves are pinched or damaged, the 
results can be numbness, interrupted neuromuscular signals, weakness, or loss of 
function.  Fortunately, Goodson (1981) determined that if the pressure was removed in a 
timely manner, the symptoms would be alleviated and the sensations in the upper 
extremity would return.   
 
Section 4: Lumbar Belt Implementation 
Hiking pack research, although less abundant than bookbag studies also provides 
recommendations about how to handle the loads placed on the body.  These include: 
placing the load closer to the body (Cook & Neumann, 1987) and utilizing a lumbar belt 
to decrease the load placed on the shoulders (LaFiandra & Harman, 2004; Mackie et at., 
2005).  Lumbar belts have been advertised to customers based on the subjective concern 
of comfort and performance.  However, it is the empirical evidence that these structures 
are able to provide for more effective load carriage that will assist in potential adaptations 
 44 
to the marching snare drum system (SDS) (Bloom & Woodhull-McNeal, 1987; Holewijn, 
1990; Reid et al., 2004).   
Holewijn (1990) recruited four college-aged males, unaccustomed to carrying 
backpacks, and placed loads of 5.4 kg and 10.4 kg inside both a military backpack that 
placed the load on the shoulders and a custom backpack that divided the load between the 
shoulders and hips through the use of a lumbar belt.  Each participant was asked to stand 
on a treadmill for 20 minutes and walk on the same treadmill at a speed of 1.3 m/s for an 
additional 20 minutes.  Data collection consisted of an unloaded, 5.4 kg military pack, 5.4 
kg custom pack, 10.4 kg military pack, and 10.4 kg custom pack.  The pressure data were 
measured at the lateral border, medial border, and center of the right where the straps 
made contact with the body.  Results indicate that placing a load on the body does cause 
an increase in skin contact pressure, particularly from the top of the shoulder along the 
back side of the strap towards the scapula.  Across all conditions, the military pack 
increased skin pressure up to 203 mmHg as compared to the custom pack (only 15 
mmHg) for both the static and dynamic conditions.  The data also showed significant 
changes between the masses being carried in the military pack; for example, standing 
while carrying the 10.4 kg load resulted in a 36% increase in pressure compared to the 
5.4 kg load.  The custom pack showed significantly lower pressure on the superior aspect 
of the shoulder than the military pack, and the military pack produced a significant 
increase in pressure on the edges of the straps within loading conditions (Holewijn, 
1990).   
Military load carriage research not only focuses on the load mass, gait 
kinematics/kinetics, and temporal variables, but forces exerted on the body as well.  
 45 
LaFiandra et al. (2004) chose to investigate the effect of a military backpack mass on the 
forces being exerted onto the individual, and how much of the total force is being exerted 
on the upper back versus lower back.  Eleven male soldiers were recruited to walk on a 
treadmill for 3 minutes at 1.34 m/s while carrying a MOLLE system loaded at 13.6 kg, 
27.2 kg, and 40.8 kg.  Two force transducers were mounted between the connections of 
the lumbar belt and the pack frame to determine if the increase in force exerted on the 
body was due to increased load mass and to determine the amount, if any, of the vertical 
forces were transmitted from the shoulders to the pelvis through the use of a lumbar belt.  
The results indicated that increasing the load mass resulted in an increase in the vertical 
forces on the lower back.  In particular, as the load mass increased from 13.6 (133 N), 
27.2 (266 N), and 40.8 kg (400 N), the mean vertical force exerted on the lower back 
increased from 41.85, 74.73, and 119.71 N, while the mean vertical force on the upper 
back/shoulders increased from 92.36, 193.10, and 282.19 N respectively.  Based on these 
findings, LaFiandra et al. (2004) determined that regardless of load mass condition, 30% 
of the vertical force generated by the pack was transmitted to the lower back with the use 
of lumbar belt, leaving 70% of the load to be managed at the shoulders.   
Mackie and colleagues (2005) also investigated the influence of load carriage, the 
use of a lumbar belt, gait speed, and shoulder strap length on shoulder strap tension and 
shoulder pressure.  Similar to Holewijn (1990), Mackie and colleagues (2005) determined 
that the load mass had the greatest influence on shoulder strap tension and shoulder 
pressure.  The researchers also found that by utilizing a lumbar belt and adjusting the 
strap length, the load could be more effectively redistributed to the hips while the 
shoulders experienced less pressure.   
 46 
Mackie et al. (2005) utilized a load carriage simulator with a programmable three 
degrees of freedom platform.  Also employed was a mannequin representing a 5th 
percentile Canadian armed forces female to replicate the characteristics of a thirteen year 
old student.  A commercially available backpack with shoulder straps and a waist belt 
was used.  The load cells were located at the top and bottom of the shoulder straps to 
ensure that desired tension levels were achieved.  Pressure sensors were also located on 
the top of the right shoulder, under the strap.  The 2 load conditions (5.3 kg =10% BM of 
mannequin and 7.9 kg =15% BM of mannequin), were placed at 5.5 cm and 11 cm 
horizontally from the inner backpack lining, high and low on the mannequin, as well as 
with and without the use of a lumbar belt.  Results of these simulated load carriage 
conditions indicated that load mass had the greatest influence on the strap forces and 
more importantly, shoulder pressures.  However, the increase in shoulder pressures was 
not proportionate as the researchers had expected.  Loads of 10% BM produced pressures 
of 222 k Pa while 15% BM produced pressures of 378 k Pa.  Utilizing a lumbar belt 
yielded a larger contact pressure for the 10% body mass condition as compared to the 
non-lumbar belt condition (246 k Pa), but decreased the pressure at 15% BM to 355 k Pa.  
Tighter shoulder straps produced an overall pressure of 350 k Pa for the 15% BM.  The 
looser straps resulted in less shoulder pressure (250 k Pa vs. 350 k Pa) and placing the 
load closer to the inside wall of the backpack also resulted in less pressure (295 k Pa vs. 
305 k Pa).  As a culmination of the results, when the lumbar belt and loose shoulder 
straps were used in conjunction, there was an overall decrease in the pressures 
experienced at the shoulder (Mackie et al., 2005). 
 47 
Martin and Hooper (2000) also investigated the effect of relocating a portion of 
the load mass to the hips through the use of five backpack belt designs on shoulder 
pressure.  The hips and pelvis can withstand greater forces (three times less sensitive to 
pressure) than other parts of the body due to the spherical head, deep acetabulum, and 
ligamentous support (Scribano, 1970).  The spherical head and acetabulum allows the 
forces to be distributed over a larger surface area, and dissipated along the curvature of 
the femoral head.  The use of these anatomical structures provided Martin and Hooper 
(2000) the basis for studying the use of a lumbar belt during load carriage. 
In total, forty-one male and thirty-one female students (no documentation was 
provided regarding previous experience with pack carriage) volunteered to carry a 20 kg 
military rucksack for 30 minutes at a speed of 3.5 km/h.  Shoulder and hip pressure was 
measured using a TekScan? F-Scan? sensor system.  In addition, the participants were 
asked to rate the comfort of each pack/belt trial.  The first belt was situated around the 
participants? waist and was made of foam.  Belt two was also made of foam, but was 
placed at the hips along with belts 3-5.  Belts 3, 4, and 5 were all made of plastic and 
various types of an air mesh material.  The results indicate that belt #1 had the largest 
maximum pressure (139 kPa), the smallest reduction in shoulder pressure (8.7%), with 
the worst comfort ratings (2.85) at the shoulders.  Belts 2, 3, 4, and 5 however, had 
maximum pressures of 82, 88, 79, and 57 kPa respectively, and were able to reduce the 
shoulder pressure to levels of 48.3, 42, 41.1, and 39.2% respectively (Martin and Hooper, 
2000).  These results provide further support that including a lumbar belt and using it 
appropriately has significant benefits to transferring a portion of the load to a more 
suitable structure such as the hips. 
 48 
Reid and colleagues (2004) investigated static load carriage and the effect of pack 
framework by implementing lateral ?stiffness? rods to a pack that already had a lumbar 
belt but no permanent frame.  The researchers investigated the forces being exerted on 
the body at the shoulders and waist to determine if the application of ?stiffness? rods 
could redistribute the load of the pack vertically, in fact shifting the load lower.  Reid et 
al. (2004) hypothesized that transferring the forces to the pelvis would minimize the 
horizontal forces in the lumbar region and would stabilize the load.  A mannequin was 
utilized and held at 10? of trunk flexion.  Reid et al. (2004) found that by attaching a rod 
from the pack to the lumbar belt, allowed the rod to become an ?active component of 
system?, transferring forces from the shoulder straps to the lumbar belt.  The rods 
reduced the vertical force applied to the upper body by 14% without a difference in shear 
forces on lumbar spine (L3/L4).  The rods also decreased the tension on the shoulders, 
and decreased the flexion moment at the lumbar (Reid et at., 2004), both adaptations that 
would presumably decrease the muscular demand at the shoulders and trunk respectively 
(Goh et al., 1998). 
In summary, the inclusion of a lumbar belt decreases the forces imposed at the 
shoulder.  Therefore, it is reasonable to conclude that the inclusion of a lumbar belt to the 
marching carrier will decrease the forces noted at the shoulder.  It is less clear as to the 
influence of the lumbar belt on gait mechanics as the pelvis will already be constrained 
by the location of the drum. 
 
 
 
 49 
Section 5: Summary 
Previous research has shown that load carriage influences COGsys excursion 
lengths, sagittal trunk movement, hip flexion, and knee flexion angle (Anderson et al., 
2007; Kinoshita, 1985; Schiffman et al., 2007).  One outcome that occurs as a result of an 
external load includes shifting the COGperson closer to the edges of the BOS (increasing 
sway) and thus compromising the balance of the individual.  In response to an external 
load being placed on the front or back of the body, individuals in a non-restricted task can 
utilize the lower extremities for shock absorption, and lowering the COGsys for increased 
stability and balance (Devroey et al., 2007; Kinoshita, 1985; Ling et al., 2004).  However 
as noted in Table 1, the marching band drummer is restricted when loaded and can?t 
always take advantage of these adaptations. 
These limitations emphasize the gap in the literature and the inability to 
automatically apply the solutions found in the load carriage literature to marching load 
carriage tasks.  The first is based on the recommendation that safe load carriage should be 
between 10 and 20% BM (Hong et al., 2008; Moore et al., 2007).  However, because the 
mass of the MDS remains constant (8.58kg) any adjustment to this load is unreasonable.  
A second constraint is that the horizontal placement of the COGdrum cannot be adjusted, 
and the rods on the carrier are not load-bearing rods similar to those found in framed 
packs.  The inability to move the COGdrum closer to the individual limits the opportunity 
to decrease the horizontal shear forces from the marching carrier at the lumbar region.  
Without load-bearing rods in the carrier, the shear forces continue to act on the lumbar 
region instead of being transferred to the rods.  A third restriction is the length of the 
body of the carrier.  The length is standard but drummers are many different heights.  The 
 50 
hiking and backpack industry suggest that the size of the pack should fit the torso (REI, 
Inc., 2010).  However, all drums must be maintained at the same level for aesthetic 
reasons.  This often precludes vertical adjustments of the load on shorter individuals as 
moving the drum too high would cause the drummers? arms to be placed in an awkward 
and unrealistic playing position.  Arm placement/location is another constraint that has an 
impact of the performance ability of marching drumline members.  Research that 
investigated rifle carriage and the restrictive arm motion that accompanies that load 
(Attwells et al., 2006; Birrell & Haslam, 2009), relates well to the restricted arm motion 
found in marching band participants.   This limited arm motion has been shown to 
increase the movement of the COGperson, leading to increases in the side-to-side motion 
which could lead to a loss of balance and potentially a fall (Birrell, Hooper, & Haslam, 
2007).  Additional kinematic changes that individuals wearing the carrier cannot employ 
are included in the following table (Table 1) including the possible implications and 
outcomes as a result of wearing an SDS. 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 51 
Table 1 
Kinematic Challenges of the SDS 
 1
Al-Khabbaz et al., 2008; 2Anderson et al., 2007; 3Attwells et al., 2006; 4Birrell & Haslam, 2009; 5Cham & 
Redfern, 2004; 6Cook & Neumann, 1987; 7Devroey et al., 2007; 8Goh et al., 1998; 9Harman et al., 1992; 
10Holewijn, 1990; 11Holt et al., 2003; 12Kinoshita, 1985; 13LaFiandra et al., 2003; 14LaFiandra et al., 2004; 
15Ling et al., 2004; 16Mackie et al., 2005; 17Martin & Hooper, 2000; 18Martin & Nelson, 1986; 19Sharpe et 
al., 2008 
 
There is much to be understood regarding the load carriage demands and 
constraints placed upon the marching snare drummer.  Including but not limited to: (a) 
determine the adaptations (postural and kinematic) in experienced marching band snare 
drummers to the load carriage and constraints of wearing a marching snare drum system: 
and (b) determine the influence of a specialized lumbar belt on these variables.  
Backpacks that utilize lumbar belts have been shown to redistribute as much as 30% of 
the vertical force from the upper back to the lower back for better load distribution and 
possibly less muscular demand on the participant (LaFiandra & Harman, 2004).  These 
findings suggest that the inclusion of a frame and lumbar belt can result in some of the 
most positive combined effects to backpack load carriage, and perhaps provide guidance 
 52 
for approaching marching load carrying tasks.  The outcome of this research will advance 
the load carriage literature and provide potential solutions to the unique population of 
marching band snare drummers.   
 
 
 
 
 
 53 
CHAPTER III 
METHODS 
 
 Thousands of marching band musicians perform demanding and strenuous 
movements each day, and a particular group of musicians perform these movements 
while carrying large cumbersome loads.   These load carriage demands were investigated 
in order to determine the influence of a SDS on the biomechanics of the marcher and to 
provide information about a population and load carriage technique that has received 
little attention in the literature.  Therefore, the purpose of this investigation was three-
fold:  1) to determine the influence of a SDS on trunk and lower extremity kinematics 
while standing and marching, 2) to determine the contact pressures and muscular demand 
placed on the erector spinae individuals wearing a SDS, and 3) to investigate the 
influence of a lumbar belt (modified SDS (MDS)) on trunk kinematics, muscular activity, 
and contact pressures.  The following chapter presents the methodology used to address 
the purposes of the present study and includes the following sections: (a) participants, (b) 
equipment, (c) procedure, and (d) statistical analysis.  The research protocol was 
approved by the Auburn University Institutional Review Board for Research Involving 
Human Subjects.   
 
 
 
 54 
Participants 
  Sixteen college-aged males (N = 12) and females (N = 4) between the ages of 19-
25 years, with drumline marching experience, specifically snare or tenor drum 
experience, were recruited from the 2010-2011 Auburn University Marching Band.  
Although there are only 9 individuals who currently participate in the marching band as 
snare drummers, other drumline members were recruited as long as they had experience 
marching at least one season (college or high school) as a snare or tenor drummer.  The 
tenor drums also place the load anteriorly and inferiorly to the drummer?s body and 
therefore have a similar load demand on the body.  All participants were clothed in a 
compression tank-top, compression shorts, and tennis shoes.  Each participant completed 
a health screening survey (Appendix A) and was excluded if a question was answered 
"yes" and it was deemed by the primary investigator that the "yes" would influence the 
participant's performance in the study.  Exclusion criteria included any current lower 
extremity, trunk, or shoulder injury or surgery, any known decrements in balance, or 
currently had an inner ear disturbance.  Each participant was asked to sign a University 
approved informed consent document to indicate his or her voluntary willingness to 
participate in this project (Appendix B). 
 
Setting 
 All testing and data collection occurred in the Sport Biomechanics Laboratory at 
Auburn University (1127 Memorial Coliseum).  The Sport Biomechanics Laboratory is a 
large enclosed laboratory with the necessary equipment to carry out this research.  
 
55 
 
Materials 
Marching Snare Drum System 
 The SDS that was used for this study and by the Auburn University Marching 
Band Drumline is comprised of the Pearl? (Pearl? Corporation, Nashville, TN, USA) 
Championship Maple FFX-1412 Snare Drum (7.67 kg) (Figure 1) with a Randall May? 
(Randall May? International, Inc., Irvine, CA, USA) Aluminum Tube Snare Carrier (0.91 
kg) (Figure 2).  The maximum length of the carrier from the peak curvature of the 
shoulder frame to the lowest bracket arm position is 78.1 cm (Figure 3), the height of the 
snare drum is 35.6 cm (Figure 3).  When worn together the total length of the SDS from 
the top of the shoulder frame to the bottom of the drum is approximately 101 cm.  The 
reason why the total length does not equal 113.7 cm is due to the 12.7 cm where the top 
of the drum sits above the bracket arm (Figure 4).  The shoulder frame is 40.64 cm in 
length, 5.1 cm in width, with a back piece that makes contact over the scapulae that is 
25.4 cm in length (Figures 5 & 6).  The abdominal plate was measured at its three widest 
parts: top (13 cm), middle (24 cm), and bottom (29 cm) (Figure 7).  
  Figure 3. Length of the carrier and height of the drum. 
 
 
78.1 
cm 
35.6 
cm 
 56 
 Figure 4. Distance between the bottom of the bracket arm and top of the drum. 
 
 
  
 
 
Figures 5 & 6. Snare drum carrier shoulder frame from the sagittal and posterior view 
respectively.  
 
5.1 
cm 
25.4 
cm 
 
40.64 
cm 
12.7 
cm 
 57 
 Figure 7. Snare drum carrier abdominal plate from a posterior view. 
 
 
Each participant used the same snare drum set, and the drum height was restricted 
to a common height based on recommendations by the percussion instructor, as would be 
the case during the regular marching season.   
 
Lumbar belt 
 The belt for this study will be referred to as lumbar belt as that is the location on 
the body upon which it will be placed.  The lumbar belt was constructed by the primary 
investigator using 2 mm neoprene (Seattle Fabrics, Inc.?, Seattle, WA, USA), 0.95 cm 
polyethylene tubing (Watts?, N. Andover, MA, USA), one roll of 1? x 60? nylon 
strapping (Sew-Ology?, Oklahoma City OK, USA), ?? closed cell foam (Seattle Fabrics, 
Inc.?, Seattle, WA, USA), and industrial strength velcro (10.1 cm x 5 cm) (Velcro USA 
Inc.?, Manchester, NH, USA).   
The initial construction consisted of the two neoprene layers were cut into a 
standard lumbar belt shape and consisted of these specifications:   
? Maximum length of the neoprene is 20? and maximum height is 8? (Figure 8a) 
? Minimum height of the neoprene is 2? tall located at the very ends of the material 
(Figure 8b) 
13 
cm 
24 
cm 
29 
cm 
 58 
? The neoprene maintained the height of 8? for the length of 16? (8? left and right 
of the center and in line with the plastic tubing), at which time was cut at a 60? 
angle (Figure 9) 
? Three pieces of neoprene material used for this belt; one solid piece and two 
smaller pieces that overlapped each other.  The largest piece had the dimensions 
described above and was the layer that faced away from the individual. (Figures 
8a & 8b)  The second and third pieces were also cut to a length of 20? but only to 
a height of 5? and 4? respectively.  This one inch difference allowed for the 
overlap between the two pieces for strength, protection of the foam, and easy 
removal of the foam in order to allow the belt to be washed. (Figures 10a & 10b) 
? The foam was cut to a length of 18? and 6? tall, following the path and angle of 
cut of the neoprene material, leaving 1? on either end for the space between the 
foam and the neoprene where the nylon strapping would exit the belt. (Figure 
10a) 
? The plastic tubing was cut to a length of 16? and placed at the top and bottom 
edges of the material just above the foam.  A pocket was then sewn at the top and 
bottom edges of the belt to encase the plastic tubing and allow for easy removal. 
(Figure 11) 
? The nylon strapping was cut into two-30? pieces.  One piece was sewn at one end 
of the belt, angled up towards the outer edge, sewn in parallel to the plastic 
tubing, angled again towards the opposite end where it exited to be used as one of 
the ends that would go around the metal tubing of the drum carrier.  This process 
 59 
was repeated with the other 30? piece with the only different being the opposite 
starting position. (Figures 8a, 8b, & 11) 
? Velcro was attached in three places for this belt-1) along the straps that exited the 
belt and were used to wrap around the drum carrier (Figure 12); 2), along the 
back of the belt to provide for attachment sites for the straps (Figures 8b, 11, & 
12); and 3) along the side facing the individual, on both sides of the neoprene 
pieces that overlapped each other for insertion and removal of the foam.  (Figures 
10a &10b) 
          
  
Figures 8a & 8b. Lumbar belt construction. Figures 8a and 8b note the overall size of the 
belt, as well as the particular stitching of the nylon straps (circle) and the placement of 
the outside velcro (single arrow). 
 Figure 9. Figure 9 notes the angle at which the corners of the belt were cut at. 
20? 
16? 
8? 6? 2? 
 60 
 
Figure 10a & 10b. Lumbar belt construction. Figure 10a notes the two separate layers of 
neoprene (circle) that were folded over each other to provide a secure closure (10b), as 
well as easy opening to insert and remove the foam (arrow).  The foam was cut to a 
length of 18?, a max height of 6?, min height of 2? at the ends, and a 60? angle taper from 
the center to the ends. 
 
 
Figure 11. Lumbar belt construction. Figure 11 notes the plastic tubing used for the belt 
design, and the pocket that was stitched for its insertion and removal.  Seen here again are 
the stitching techniques for securing the nylon strapping and the velcro attachment. 
 
 Figure 12. Lumbar belt construction. Figure 12 notes the velcro attachment to the straps 
(orange arrow) seen exiting the belt (circle), as well as the back of the belt attachment 
(white arrow) to which the straps will connect to for a secure hold of the drum to the 
individual. 
 
 
 61 
Instrumentation 
Kinematics 
 Three dimensional trajectories of the foot, ankle, knee, hip, and trunk were 
obtained with a ten camera Vicon? MX motion analysis system (Vicon?, Los Angeles, 
CA, USA) with a sampling frequency of 288 Hz (Devroey et al., 2007).  To track the 
movement strategies of the lower extremities and trunk, 34 retroreflective markers were 
attached to each participant?s body (Table 2 and Figures 13-16).  The markers used 
were spherical 14mm retro-reflective markers (MKR-6.4?, B&L Engineering?, Tustin 
CA, USA) attached with double-sided tape (CLEAR-1R36?, Hair Direct Inc.?, 
Bainbridge, PA, USA). The standard Plug-In-Gait? marker placement protocol (Vicon?) 
was used with the sacrum marker being used in place of a left and right PSIS in order to 
accommodate the drum carrier and lumbar belt placements.    
Table 2 
Names and positions of markers used in Plug-In-Gait? Model 
 
 62 
                
Figures 13 & 14. Anterior and posterior views of the Plug-In-Gait marker locations. 
 
 
                Figures 15 & 16. Sagittal views of the Plug-In-Gait? marker locations. 
 
 63 
Joint angles were calculated by Vicon? software using a modified Plug-In-Gait 
model (Devroey et al., 2007) (Figures 17-20).   
 
 
 
     
 
 
      
 
 
Figures 17 & 18. Trunk and hip flexion/extension with flexion positive and extension 
negative relative to the absolute angle (Attwells et al., 2006). 
 
                                     
Figures 19 & 20. Knee flexion/extension with flexion positive and extension negative.  
Ankle plantarflexion and dorsiflexion with dorsiflexion positive and plantarflexion 
negative (Attwells et al., 2006). 
 
 
  ANK 
 TOE 
KNE 
 
 SLDR 
 RASI 
ASIS 
KNE 
 64 
 
Ten cameras were positioned to record all three cardinal planes (frontal, sagittal, 
and transverse) of the participant during one full stride, defined as right heel contact to 
right heel contact (Figure 21). 
 
 
Figure 21. Vicon motion capture camera setup. 
 The gait cycle used for the current study was defined as right foot contact, right 
foot midstance, right foot heel-off, left foot contact, right foot toe-off, left foot midstance, 
left foot heel-off, and a second right foot contact (J. Perry, 1992) (Figure 22).  This cycle 
was used because it incorporates both phases of a gait cycle (stance and swing).  Both the 
right and the left leg are needed and consequently included for one complete stride, and 
therefore, periods during the gait cycle can be evaluated.  The participants were asked to 
walk utilizing a specific marching gait (glide step) in the laboratory instead of on the 
marching field or road (parade comparison) at a set frequency (144 beats/min) that most 
resembles performance tempos.   
 65 
   
Figure 22. Functional phases of gait (Kirtley, 2008) 
 
Surface Electromyography (sEMG) 
Muscle activity was collected using 1) Noraxon? MyoSystem 1200 tethered 
surface EMG receiver (Noraxon? U.S.A. Inc., Scottsdale, AZ, USA) (Figure 23), and 2) 
.Noraxon? Telemyo 2400R-World Wide Telemetry receiver (Noraxon? U.S.A. Inc., 
Scottsdale, AZ, USA) (Figure 24a), connected via preamplifier leads to a Noraxon? 
Telemyo 2400T V2 wireless transmitter (Noraxon? U.S.A. Inc., Scottsdale, AZ, USA) 
(Figure 24b).  The electrodes were Trace-Rite? Ag/AgCl ECG monitoring surface 
electrodes (Trace-Rite?, Bio-Detek Inc., Pawtucket, RI, USA) (Figure 25).  The software 
utilized for data collection and processing was the MyoResearch XP Master Edition? 
1.08.15 (Noraxon? U.S.A. Inc., Scottsdale, AZ, USA).  The EMG signal was sampled at 
1000Hz, and post processing included full wave rectification and a finite impulse 
response (FIR) filter (gain set at 1000, bandpass of 10-350Hz, and a notch filer of 59.5-
60.5Hz).  For the standing condition, the mean linear envelope was calculated for 1s.  For 
the gait condition, the mean linear envelope of the left and right muscles were obtained 
for one stride length (Devroey et al., 2007).  
 66 
 
Figures 23. Noraxon MyoSystem? 1200 Tethered Surface EMG receiver 
 
 
          Figure 24a & 24b. Noraxon Telemyo? 2400R ?World Wide Telemetry Receiver & 
Noraxon? Telemyo 2400T V2 wireless transmitter 
 
 
Figure 25. Erector spinae sEMG electrode placement 
 
 
 
 67 
Drum Carrier Pressure  
 Pressure was measured using a Tekscan? I-Scan System? (v 6.10.1) pressure 
distribution measurement system (Tekscan?, Inc., South Boston, MA, USA) (Figure 26).  
This Tekscan? system consists of three separate pressure sensors, allowing the researcher 
to collect data simultaneously from the superior aspect of both shoulders (Sensor #9801) 
and well as the lower abdominal/pelvic girdle area (Sensor #9830) (Figures 27 & 28).  
The software utilized for data collection was the Tekscan? I-Scan? Research version 
6.10.1 with Sensor Map #9801 and #9830 enabled (Tekscan?, Inc., South Boston, MA, 
USA).  Both average and peak contact pressures were measured in the areas between the 
carrier and the participant, specifically both shoulders, and the abdominal/pelvic region.  
Average pressure was calculated by the total force exerted on the sensor divided by the 
number of individual cells that are being triggered in a specific body region.  Peak 
pressure refers to the highest pressure recorded for a given pressure sensel (Bossi et al., 
2000).  Before each trial, the sensors were calibrated to the load (8.58 kg) of the drum 
system, and placed on the inside of the carrier's shoulder regions (Figure 5) and 
abdominal plate (Figure 7).  The placement of the sensors (row and column designation) 
were marked with chalk onto the compression tank-top (Martin and Hooper, 2000).  The 
I-Scan? Research v. 6.10.1 system sampled/scanned at 160 Hz.   
 
 68 
 
Figure 26. Tekscan? I-Scan System? (v 6.10.1) 
 
 
Figure 27. Tekscan? I-Scan System? (v 6.10.1) Pressure Sensor #9801 
 
 
 
Figure 28. Tekscan? I-Scan System? (v 6.10.1) Pressure Sensor #9830 
 
The sEMG and joint/gait kinematics were synchronized with a Vicon? MX 
Control 64-channel A/D board (Vicon?, Los Angeles, CA, USA) and all trials were 
videotaped with a Canon? 3CCD Digital Video Camcorder GL2 NTSC (Canon? U.S.A., 
Inc., Lake Success, NY, USA) for motion verification.  Kinematics were analyzed with 
the Nexus Polygon Software? (Vicon?, Los Angeles, CA, USA).  All data were saved on 
a computer that is stored and locked inside the Sport Biomechanics Laboratory for post 
processing.   
 
 
 69 
Design and Procedures 
 The participants reported to the Sport Biomechanics Laboratory for one day of 
testing.  After arrival to the lab, each participant was given the Auburn University 
Institutional Review Board approved informed consent document and preliminary 
medical questionnaire (Appendix A).  By signing the informed consent, participants 
indicated voluntary participation in the project.  The medical questionnaire was used as a 
screening device to eliminate anyone who had 1) sustained a lower extremity, trunk, or 
shoulder injury within the last six months, 2) an inner ear disorder or balance decrements, 
or 3) an allergy to adhesives.   
After meeting the criteria for participation, both the mass and height of the 
individuals were measured for normalizing purposes.  The participants were then asked to 
change into spandex shorts and tank-tops to minimize clothing artifact and disrupting the 
pressure sensor readings.  Anthropometric information was collected and entered in the 
Vicon? software.   
The skin was prepared for surface electromyography electrode placement 
following the recommendations of the Surface Electromyography for the Non-Invasive 
Assessment of Muscles (SENIAM) (Hermens et al., 2000).  The skin area was abraded to 
reduce the electrical impedance of the skin and cleaned with a 70% alcohol solution.  The 
electrodes were placed at a 2cm inter-electrode spacing, 3cm lateral to the midline 
(Anderson et al., 2007; Seroussi & Pope, 1987).  The erector spinae (ES) at the L1 level 
was identified by palpation and manual muscle testing as the participant performed trunk 
extension from a supine position (Figure 25).  Electrodes were placed on the muscle 
belly parallel to the direction of the muscle fibers.  A ground electrode was placed on the 
 70 
bony aspect of the C7 vertebrae.  Maximum voluntary isometric contraction-used for 
normalization (MVIC) was then collected.  
Thirty four trial markers were affixed to the spandex clothing and to the skin with 
double-sided tape and placed according to the requirements of the Vicon? Plug-In-Gait? 
model with slight modifications to accommodate the drum carrier and lumbar belt.  A 
static capture was performed while the participants stood in the center of the capture 
volume in an anatomical T-pose position with shoulders abducted and hips at shoulder 
width apart.  The markers were identified and labeled to comply with the Vicon? Plug-In-
Gait? model modified for this study.   
The pressure sensors were taped underneath the shoulder carrier and abdominal 
plate of the drum carrier at the beginning of each participant's data collection session.  
Two pressure sensors (#9801) were placed directly over the location where the 
SDS/MDS made contact with the body: over the superior aspect of the shoulders, angled 
towards the midline, crossing over the clavicle lateral from the sternum (Martin and 
Hooper, 2000) (Figures 29 & 30).  The third sensor (#9830) was placed across the lower 
abdominal/pelvic region, directly over the location where the SDS/MDS makes contact 
with the body.  The midline of the sensor was placed at the midline of the body (Figures 
29 & 30).  
 
 71 
   
Figures 29-30. Tekscan I-Scan System? Pressure Sensors #9801 (shoulders) and #9830 
(abdominal/pelvic area) with reference to the carrier placement respectively. 
 
 
Procedures 
These procedures were consistent throughout all three loading conditions (unloaded, 
SDS, MDS), beginning with the unloaded (U) condition followed by a random 
assignment to the SDS or MDS loading conditions.   The only difference that should be 
noted is the inclusion of contact pressure measurements during the SDS and MDS 
conditions that were taken before and after the stationary trials. 
1) Participants stood for one minute while sEMG and joint kinematics were 
collected during last five seconds.  While standing, the participants played a 
warm-up exercise on the drum that is a standard amongst drumlines called "8 on a 
hand".   
2) Participants then marched through the capture volume (glide step gait) while 
sEMG, and joint/gait kinematic data were collected.  The participants continued 
to play the warm-up exercise on the drum throughout the data collection.  
Participants were asked to march a 57.15cm step size (standard "8-to-5" marching 
step), and the step frequency was set at 144 beats/minute. 
 72 
3) Participants then performed a marching task for one minute in the gymnasium 
located adjacent to the Sport Biomechanics Laboratory.   
a. The participants marched in a straight line incorporating the glide step gait 
and a mark time (marching in place) phase, while playing the warm-up 
exercise on the drum.  Participants were asked to march a 57.15cm step 
size (standard "8-to-5" marching step), and the step frequency was set at 
144 beats/minute. 
4) Following the marching task, the participants reentered the laboratory and 
marched through the capture volume (glide step gait) while sEMG and joint/gait 
kinematic data were collected.  The participants continued to play the warm-up 
exercise on the drum throughout the data collection.  Participants were asked to 
march a 57.15cm step size (standard "8-to-5" marching step), and the step 
frequency was set at 144 beats/minute. 
5) Participants were then asked stand in a stationary position again for one minute 
while surface EMG and joint kinematics were captured during the last five 
seconds.   
 
Pressure data, indicated above, was collected before and after the last stationary trial 
during the SDS and MDS conditions.  Contact pressure procedures consisted of the 
participants standing for one minute while playing the standard "8 on a hand" warm-up 
exercise.  Data were collected during the first 5 seconds after being loaded and at 55 
seconds before the first stationary trial, and again at 55 seconds after the last stationary 
trial. 
 73 
During the stationary trials, the participants played the same warm-up exercise on the 
drum as during the dynamic trial, with the feet externally rotated to face a 45? angle.  
During the unloaded condition, the arms were in the playing position to control for arm 
placement throughout the conditions, as well as to mimic the requirements that occur 
during practice of ?air playing? when learning new music. There was a period of at least 
5 minutes of rest in between each loading condition.  
 
Experimental Design & Data Analysis 
 All statistical analyses were conducted using SPSS software (version 17.0; SPSS 
Inc., Chicago, IL, USA) and an alpha level of statistical significance was set apriori at p 
? 0.05 (?SPSS for Windows, 2007?); all data were stored in an Excel database and 
imported into SPSS for analysis.  During stationary trials, the data for sEMG and joint 
kinematics was analyzed for one second during the last five seconds of a minute trial.  
For contact pressure, the data were analyzed during the initial five seconds of a one 
minute test (Pre5), the last five seconds of a minute test (Pre55), and the last five seconds 
of a minute test after the marching exercise (Post55).   
To investigate the effects of load carriage on EMGAVG and then EMGPEAK, the 
data were analyzed by conducting a 2 (movement) x 3 (loading condition) x 2 (time) 
repeated measures ANOVA.  The independent variable of movement had two levels: 
stationary and dynamic.  The independent variable of load had three levels: unloaded 
(body mass-BM), Standard snare drum system (SDS), and modified snare drum system 
(MDS).  The independent variable of time had two levels: pre and post.   
 74 
To investigate the effects of load carriage on contact pressure (kPa) at the 
shoulders and then the abdominal/pelvic region, the data were analyzed by conducting a 2 
(loading condition) x 2 (body placement) x 3 (time) repeated measures ANOVA.  The 
independent variable of load had two levels: SDS and MDS.  The independent variable of 
body placement had two levels: shoulders and abdominal/pelvic region.  The independent 
variable of time had three levels: Pre5, Pre55, and Post55.   
To investigate the effects of load carriage on joint kinematics, the data were 
analyzed by conducting a 3 (loading condition) x 2 (movement) x 2 (time) repeated 
measures MANOVA.  The independent variable of load had three levels: unloaded (body 
mass-U), Standard snare drum system (SDS), and modified snare drum system (MDS).  
The independent variable of movement had two levels: stationary and dynamic.  The 
independent variable of time had two levels: pre and post.  The dependent variables were 
trunk angle (TAAVG), hip angle (HAAVG), knee angle (KAAVG), and ankle angle (AAAVG). 
Knee angle (KAAVG) and ankle angle (AAAVG) were also investigated during the 
foot flat phase of the gait cycle.  Two 3 (loading condition) x 2 (time) repeated measures 
ANOVAs were conducted with the same independent and dependent variables discussed 
above.  
For all statistical analyses, post hoc repeated measures ANOVAs were performed 
for each dependent variable of the multivariate repeated measures ANOVA that 
demonstrated statistical significance.  Follow-up pairwise comparisons were then 
performed for significant findings from the ANOVA analyses to determine statistical 
significance between the levels of the independent variables.  
 75 
CHAPTER IV 
 
RESULTS 
 
 
 This study was designed to determine the influence of marching band snare-drum 
carriers on marching percussionists, who often complain of shoulder and low back 
discomfort.    Specifically, the purpose of this investigation was addressed in three 
questions: 1) to determine the influence of a standard marching snare drum system (SDS) 
and a modified drum system (MDS) on the trunk and lower extremity kinematics while 
standing and playing (stationary) and marching forward and playing (dynamic), 2) to 
determine the muscular demands on the erector spinae (ES) of individuals wearing an 
SDS and a MDS, and 3) to determine the contact pressures at the shoulders and 
abdominal region of individuals wearing an SDS and a MDS.  The following chapter 
presents the results of the current study and includes the following sections: Section 1: 
Participant demographics, Section 2: The effect of the SDS and MDS systems on joint 
kinematics, Section 3: The effect of the SDS and MDS systems on ES muscle activity, 
and Section 4: Difference in contact pressures between the SDS and MDS loading 
conditions, at the shoulder and abdominal regions.  
 
 
 
 
 76 
Section 1: Participant Demographics 
 Sixteen members (12 males, 4 females) of the Auburn University Marching Band 
Drumline volunteered for participation in the current study and met the initial 
qualifications.  Upon answering the medical questionnaire, all volunteers were included 
in the study and each participant signed an Informed Consent document approved by the 
University Institutional Review Board for Human Subjects (Appendix B).  The averages 
of the participants? demographics are summarized in Table 3. 
Table 3 
Participant Demographics 
    Mean 
Standard 
Deviation 
Age (years)   20.81 1.22 
Height (m) 
 
1.79 8.17 
Mass (kg)  81.64 23.07 
Years Marched 
 
6 2.28 
Years Marched Snare/Tenor  3.63 2.13 
Drum Mass to BM (%)   11.11 2.35 
 
 
 
Section 2: Joint Kinematics 
 
 In an attempt to understand the demands placed upon the human body when one 
wears a SDS, the angular kinematics of the trunk, hip, knee and ankle were investigated.  
These data were then compared to an unloaded condition and a loading configuration that 
included a lumbar belt added to the SDS, which created the MDS condition.  Next, the 
angular kinematics were investigated during a stationary condition as well as during a 
marching forward condition.   
 In order to determine significance, a 3 (condition: unloaded (U), SDS, MDS) x 2 
(movement: stationary, dynamic) x 2 (time: pre, post) multivariate repeated measures 
(MANOVA) across all four joint angles (Trunk Angle-TAAVG, Hip Angle-HAAVG, Knee 
 77 
Ankle-KAAVG, & Ankle Angle-AAAVG) with a p value set apriori at p < 0.05 was 
employed and yielded a significant interaction effect for loading condition (U, SDS, and 
MDS) and movement (stationary versus dynamic) on the kinematics of trunk angle 
(TAAVG), hip angle (HAAVG), and ankle angle (AAAVG), (Wilks?s ? = 0.100, F(2,14) = 
10. 297, p < 0 .001, ?2 = 0.684, Power = 1.000).  This indicates that different loading 
conditions altered joint kinematics differently across movement conditions.  Descriptive 
statistics for all dependent variables are presented in Appendix C. 
 
Average Trunk Angle Over One Stride  
The MANOVA revealed that there was a significant interaction between loading 
condition and movement condition.  As a result of this interaction, a follow-up repeated 
measures ANOVA for TAAVG during the stationary trials for all three loading conditions 
was then conducted.  A significant effect of condition on TAAVG was found (F(2,14) = 
13.929, p < 0.001, ?2 = 0.499,  Power = 0.996).  Bonferroni post hoc tests yielded a 
significant difference between the U (18.70?) and SDS (21.82?) conditions (p = 0.004) 
and the U (18.70?) and MDS (21.94?) conditions (p = 0.009).  Both the SDS and MDS 
conditions yielded a larger TAAVG, indicating that the participants were in a more flexed 
position during those two conditions, than during the unloaded condition.  Although a 
significant effect of condition for TAAVG was found (F(2,14) = 13.929, p < 0.001, ?2 = 
0.499,  Power = 0.996), there was no significant difference between the SDS (21.82?) and 
MDS (21.94?) conditions (p = 0.824), but the MDS condition did yield a larger TAAVG 
which indicates that the participants were in a more flexed position during that condition 
(Figure 31).    
 78 
Figure 31.  Effects of loading conditions on TAAVG during stationary trials.                    
U-SDS: *p = 0.004. U-MDS: +p = 0.009.  
 
A follow-up repeated measures ANOVA for TAAVG during the dynamic trials for 
all three loading conditions was conducted.  A significant effect of loading condition on 
TAAVG was found (F(2,14) = 88.236, p < 0.001, ?2 = 0.855,  Power = 1.000).  Bonferroni 
post hoc tests yielded a significant difference specifically between the U (15.79?) and 
SDS (22.96?) conditions (p < 0.001), and the U (15.79?) and MDS (22.71?) conditions (p 
< 0.001).  Both the SDS and MDS conditions yielded a larger TAAVG, indicating that the 
participants were experiencing more trunk flexion during those conditions than during the 
unloaded condition.  Although an overall significant effect of loading condition on 
TAAVG was found (F(2,14) = 88.236, p < 0.001, ?2 = 0.855,  Power = 1.000), there was 
no significant difference between the SDS (22.96?) and MDS (22.71?) conditions (p = 
1.000), it should be noted that the SDS condition did yield a larger TAAVG which 
0
5
10
15
20
25
30
Unloaded SDS MDS
An
gle
 (d
eg
ree
s) 
Loading Conditions 
Effects of Two Loading Systems on TAAVG 
During Stationary Trials 
* 
* + 
+ 
 79 
indicates that the participants were in a more flexed position during that condition 
(Figure 32).    
 Figure 32.  Effects of loading conditions on TA
AVG during dynamic trials.                      
U-SDS, U-MDS: *+p < 0.001. 
 
 
Average Hip Angle Over One Stride  
  
The findings from the repeated measures MANOVA revealed that there was a 
significant interaction between loading condition and whether or not the participants were 
stationary or dynamic.  As a result of this interaction, a follow-up repeated measures 
ANOVA for HAAVG during the stationary trials for all three loading conditions was 
conducted.  A significant effect of loading condition on HAAVG found (F(2,14) = 4.440, p 
= 0.020, ?2 = 0.228,  Power = 0.719).  Bonferroni post hoc tests yielded a significant 
difference only between the U (73.82?) and SDS (77.08?) loading conditions (p = 0.043).  
There was no significant difference between the U (73.82?) and MDS (78.09?) loading 
conditions (p = 0.057), or between the SDS (77.08?) and MDS (78.09?) conditions (p = 
0
5
10
15
20
25
30
Unloaded SDS MDS
An
gle
 (Deg
ree
s) 
Loading Conditions 
Effect of Two Loading Systems on TAAVG 
During Dynamic Trials 
* 
+ 
+ 
* 
 80 
1.000).  These findings indicate that SDS produced a more vertical thigh position than the 
unloaded condition, but the MDS condition, though not significant, yielded the most 
vertical thigh position (Figure 33). 
 Figure 33.  Effects of loading conditions on HA
AVG during stationary trials.                    
U-SDS: *p = 0.043. 
 
A follow-up repeated measures ANOVA for HAAVG during the dynamic trials for 
all three loading conditions was conducted.  A significant effect of condition on HAAVG 
was found (F(2,14) = 4.638, p = 0.018, ?2 = 0.236,  Power = 0.739).  Bonferroni post hoc 
tests however did not yield a significant difference between any of the loading 
conditions: U (86.48?) and SDS (88.30?) conditions (p = 0.097), U (86.48?) and MDS 
(88.72?) conditions (p = 0.111), and SDS (88.30?) and MDS (88.72?) conditions (p = 
1.000).  Similar to the results found during the stationary movement condition, although 
there was no significant difference between the three loading conditions, the MDS 
condition yielded a larger HAAVG, indicating less hip flexion angle (Figure 34). 
60
65
70
75
80
85
90
Unloaded SDS MDS
An
gle
 (Deg
ree
s) 
Loading Conditions 
Effects Two Loading Systems on HAAVG 
During Stationary Trials 
* 
* 
 81 
 
Figure 34.  Effects of loading conditions on HAAVG during dynamic trials.                       
U-SDS: p = 0.097.  U-MDS: p = 0.111. SDS-MDS: p = 1.000.  
 
Average Ankle Angle Over One Stride 
 
The findings from the repeated measures MANOVA revealed that there was a 
significant interaction between loading condition and movement condition (i.e., 
stationary or dynamic).  As a result of this interaction, a follow-up repeated measures 
ANOVA for AAAVG during the stationary trials for all three loading conditions was 
conducted.  A significant effect of loading condition on AAAVG was found (F(2,14) = 
32.516, p < 0.001, ?2 = 0.684,  Power = 1.000).  Post hoc tests using Bonferroni yielded a 
significant difference specifically between the U (91.35?) and SDS (93.75?) loading 
conditions (p < 0.001), and between the U (91.35?) and MDS (93.76?) loading conditions 
(p < 0.001).  In both cases respectively, the SDS and MDS conditions yielded a larger 
AAAVG, indicating that the ankle was in a more plantarflexed position during those 
conditions in comparison to the unloaded condition (Figure 35).   
78
80
82
84
86
88
90
92
94
Unloaded SDS MDS
An
gle
 (Deg
ree
s) 
Loading Conditions 
Effects of Two Loading Systems on HAAVG 
During Dynamic Trials 
 82 
 
Figure 35.  Effects of loading conditions on AAAVG during stationary trials.                   
U-SDS, U-MDS: *+p < 0.001. 
 
A follow-up repeated measures ANOVA for AAAVG during the dynamic trials for 
all three loading conditions was conducted.  There was no significant effect of loading 
condition on AAAVG found (F(2,14) = 0.763, p = 0.475, ?2 = 0.048,  Power = 0.167) 
between the U (90.35?) and SDS (90.51?) loading conditions, between the U (90.35?) and 
MDS (89.80?) loading conditions, or  between the SDS (90.51?) and MDS (89.80?) 
loading conditions.  All three conditions yielded ankle angles very close to the 
regimented 90? required during a marching performance (Figure 36).   
82
84
86
88
90
92
94
96
98
Unloaded SDS MDS
An
gle
 (Deg
ree
s) 
Loading Conditions 
Effects of Two Loading Systems on AAAVG 
During Stationary Trials 
+ 
+ 
* 
* 
 83 
 
Figure 36.  Effects of loading conditions on AAAVG during dynamic trials.                       
U-SDS: p = 1.000.  U-MDS: p = 0.923. SDS-MDS: p = 0.579.  
 
Average Knee Angle During Foot Flat (FF) 
 
In addition to investigating the effects of the SDS and MDS loading conditions on 
the joint kinematics of the body during a stride, the author also investigated the average 
joint kinematics of the knee during the FF phase of the stride length (KAAFF).  In order to 
determine significance, a 3 (condition: unloaded (U), SDS, MDS) x 2 (time: pre, post) 
multivariate repeated measures (ANOVA) across knee ankle with a p value set apriori at 
p < 0.05 was employed (Wilks?s ? = 0.928, F(2,14) = 0.545, p = 0 .592, ?2 = 0.072, 
Power = 0.122).  The multivariate repeated measures ANOVA did not yield a significant 
effect for loading condition, (F(2,14) = 1.114, p = 0.341, ?2 = 0.069, Power = 0.227), nor 
time (pre versus post) on the KAAFF, (F(2,14) = 0.125, p = 0.729, ?2 = 0.008, Power = 
0.063): U Pre (176.45?), SDS Pre (175.8?), MDS Pre (175.17?), U Post (176.21?),  SDS 
Post (175.54?), & MDS Post (175.47?).  This indicates that the loading conditions do not 
82
84
86
88
90
92
94
96
Unloaded SDS MDS
An
gle
 (De
gre
es)
 
Loading Conditions 
Effects of Two Loading Systems on AAAVG 
During Dynamic Trials 
 84 
alter joint kinematics differently during the foot flat phase of the gait cycle.  Descriptive 
statistics for all dependent variables are presented in Appendix D (Figure 37). 
 Figure 37.  Effects of loading conditions on KA
AFF.     pLoad= 0.341. pTime= 0.729.  
 
 
Average Ankle Angle During Foot Flat (FF) 
 
In addition to investigating the effects of the SDS and MDS loading conditions on 
the joint kinematics of the body during a stride, the author also investigated the average 
joint kinematics of the ankle during the FF phase of the gait cycle (AAAFF).  In order to 
determine significance, a 3 (condition: unloaded (U), SDS, MDS) x 2 (time: pre, post) 
multivariate repeated measures (ANOVA) across knee ankle with a p value set apriori at 
p < 0.05 was employed (Wilks?s ? = 0.917, F(2,14) = 0.630, p = 0 .547, ?2 = 0.083, 
Power = 0.135). The multivariate repeated measures ANOVA did not yield a significant 
effect for loading condition (modified, SDS and MDS), (F(2,14) = 0.467, p = 0.631, ?2 = 
0.030, Power = 0.119), nor time (pre versus post) on the AAAFF, (F(2,14) = 0.691, p = 
0.419, ?2 = 0.044, Power = 0.122): U Pre (87.92?), SDS Pre (87.51?), MDS Pre (87.53?), 
174.5
175
175.5
176
176.5
177
Pre Post
An
gle
 (Deg
ree
s) 
Loading Conditions 
Effects of Loading Conditions on KA During 
the FF Phase of the Gait Cycle 
U
SDS
MDS
 85 
U Post (88.48?),  SDS Post (88.56?), & MDS Post (87.40?).  Similar to the kinematic 
results at the knee, this analysis indicates that loading conditions do not alter joint 
kinematics differently during the foot flat phase of the gait cycle.  Descriptive statistics 
for all dependent variables are presented in Appendix D (Figure 38). 
 
 
Figure 38.  Effects of loading conditions on AAAFF.  pLoad = 0.631. pTime = 0.419 
 
 
Section 3: Erector Spinae Muscle Activity 
 
In addition to examining the effects of an SDS and MDS system on joint 
kinematics, this research study also investigated the outcome of both the SDS and MDS 
systems on the muscular demands of the erector spinae muscle group (ES) while standing 
and playing in place (stationary) and marching forward while playing (dynamic).  This 
section includes the results from the SDS and MDS loading conditions compared to the 
unloaded condition in terms of average (EMGAVG ) and peak EMG (EMGPEAK ) levels, 
including follow-up tests for specific conditions. 
86
86.5
87
87.5
88
88.5
89
89.5
90
Pre Post
An
gle
 (Deg
ree
s) 
Loading Conditions 
Effects of Loading Conditions on AA During 
the FF Phase of the Gait Cycle 
U
SDS
MDS
 86 
Average sEMG 
A 2 (movement: stationary vs. dynamic) x 3 (loading condition: U, SDS and 
MDS) x 2(time: pre and post a one minute marching exercise) multivariate repeated 
measures ANOVA with a p value set apriori at p < 0.05 was employed and yielded a 
significant interaction effect of loading condition and time on the average muscle activity 
of the ES, (F(2,14) = 1.006, p = 0.041, ?2 = 0.252, Power = 0.618).  This indicates that 
the muscle activity of the ES differed across the loading conditions, pre and post the one 
minute marching exercise.  Descriptive statistics for all dependent variables are presented 
in Appendix E. 
As a result of this interaction, a follow-up repeated measures ANOVA for 
EMGAVG during the pre-trials for all three loading conditions was conducted, and a 
significant effect of loading condition on EMGAVG was found (F(2,14) = 28.958, p < 
0.001, ?2 = 0.674,  Power = 1.000).  Post hoc tests using Bonferroni yielded a significant 
difference between the U (1.70% MVIC) and SDS (3.57% MVIC) loading conditions (p 
< 0.001), and MDS (3.64% MVIC)  (p < 0.001) loading condition with the SDS and 
MDS conditions yielding higher EMGAVG levels than the unloaded condition.  There was 
no significant difference between the SDS (3.57% MVIC) and MDS (3.64% MVIC) 
conditions (p = 1.000).  This indicates that the participants' average ES muscle activity 
increased while wearing  the SDS and MDS systems during both stationary and dynamic 
tasks, before participating in the one minute marching exercise (Figure 39).    
 
 87 
 
Figure 39.  Effects of loading conditions on ES EMGAVG before a one minute marching 
exercise.  U-SDS, U-MDS: *+p < 0.001. 
 
 
A follow-up repeated measures ANOVA for EMGAVG during the post trials for all 
loading conditions across the non-included variable of movement (stationary vs. 
dynamic) was conducted.  A significant effect of loading condition on EMGAVG was 
found (F(2,14) = 29.426, p < 0.001, ?2 = 0.710,  Power = 1.000).  Post hoc tests using 
Bonferroni yielded a significant difference between the U (1.72% MVIC) and SDS 
(4.01% MVIC) loading conditions (p < 0.001), and MDS (3.68% MVIC) with the SDS 
and MDS conditions yielding higher EMGAVG level.  There was no significant difference 
between the SDS (4.01% MVIC) and MDS (3.69% MVIC) conditions (p = 0.558).  This 
indicates that the participants' average ES muscle activity was increased while wearing 
the SDS and MDS systems during both stationary and dynamic tasks, after participating 
in the one minute marching exercise (Figure 40).    
 
0
1
2
3
4
5
6
Unloaded SDS MDS
% M
VIC sE
MG 
(%
) 
Loading Conditions 
Effects of Loading Conditions on ES EMGAVG  
Before a One Minute Marching Exercise 
* + 
* + 
 88 
 
Figure 40.  Effects of loading conditions on ES EMGAVG after a one minute marching 
exercise.  U-SDS: *p < 0.001. 
 
 
Peak sEMG 
 
A 2 (movement: stationary vs. dynamic) x 3 (loading condition: U, SDS and 
MDS) x 2(time: pre and post marching) a multivariate repeated measures ANOVA with a 
p value set apriori at p < 0.05 yielded a significant interaction effect of loading condition 
and movement on the peak muscle activity of the ES, (F(2,14) = 6.996, p = 0.005, ?2 = 
0.412, Power = 0.883).  This indicates that the loading conditions altered the muscle 
activity of the ES differently across movement conditions.  Descriptive statistics for all 
dependent variables are presented in Appendix E. 
The findings from the repeated measures ANOVA revealed that there was a 
significant interaction between loading condition and whether or not the participants were 
stationary or dynamic.  As a result of this interaction, a follow-up repeated measures 
ANOVA for EMGPEAK during the stationary trials for all loading conditions was 
0
1
2
3
4
5
6
7
Unloaded SDS MDS
%M
VIC sE
MG 
(%
) 
Loading Conditions 
Effects of Loading Conditions on ES EMGAVG 
After a One Minute Marching Exercise 
* 
* 
+ 
+ 
 89 
conducted.  A significant effect of loading condition on EMGPEAK was found (F(2,14) = 
25.847, p < 0.001, ?2 = 0.665,  Power = 1.000).  Post hoc tests using Bonferroni yielded a 
significant difference between the U (3.05% MVIC) and SDS (5.61% MVIC) loading 
conditions (p < 0.001), and the U (3.05% MVIC) and the MDS (5.30% MVIC) loading 
conditions (p = 0.001), with both the SDS and MDS conditions yielding higher EMGPEAK 
levels.  There was no significant difference between the SDS (5.61% MVIC) and MDS 
(5.30% MVIC) conditions (p = 0.995).  This indicates that the participants' peak ES 
muscle activity increased and was affected when wearing the SDS and MDS systems 
during the stationary task (Figure 41).    
 
 Figure 41.  Effects of loading conditions on ES EMG
PEAK during stationary trials.          
U-SDS: *p < 0.001. 
 
 
A follow-up repeated measures ANOVA for EMGPEAK during the dynamic trials 
for all loading conditions was conducted.  A significant effect of loading condition on 
0
1
2
3
4
5
6
7
8
9
Unloaded SDS MDS
% M
VIC sE
MG 
(%
) 
Loading Conditions 
Effects of Loading Conditions on ES EMGPEAK 
During Stationary Trials 
* 
* 
+ 
+ 
 90 
EMGPEAK was found (F(2,14) = 20.834, p < 0.001, ?2 = 0.654,  Power = 1.000).  Post hoc 
tests using Bonferroni yielded a significant difference between the U (5.62% MVIC) 
loading condition and SDS (9.74% MVIC) (p = 0.005) and the U (5.62% MVIC) and 
MDS (10.25% MVIC) conditions (p < 0.001), with both the SDS and MDS conditions 
yielding higher EMGPEAK levels.  There was no significant difference between the SDS 
(9.74% MVIC) and MDS (10.25% MVIC) conditions (p = 1.000).  This indicates that the 
participants' peak ES muscle activity increased and was affected when wearing both the 
SDS and MDS systems during the dynamic task (Figure 42).    
 
 
Figure 42.  Effects of loading conditions on ES EMGPEAK during dynamic trials.            
U-SDS: *p = 0.006. U-MDS: +p < 0.001. 
 
 
 
Section 4: Contact Pressure 
 
In addition to kinematics and EMG, contact pressure at the abdominal area and 
shoulders were also examined.  Contact pressures were investigated while in a stationary 
0
2
4
6
8
10
12
14
16
18
Unloaded SDS MDS
% M
VIC sE
MG 
(%
) 
Loading Conditions 
Effects of Loading Conditions on ES EMGPEAK 
During Dynamic Trials 
* 
* 
+ 
+ 
 91 
position for one minute while playing, with and without the presence of the lumbar belt 
(SDS vs. MDS), before and after a one minute marching exercise.  These measures were 
taken at three times during the data collection: Pre5-during the initial 5 seconds of being 
loaded before the marching exercise, Pre55-during the last 5 seconds of the initial loading 
before the marching exercise, and Post55-during the last 5 seconds of marching in place 
after the one minute exercise.      
 A 2 (shoulder: right and left) X 3 (time: pre5, pre55, and post55) repeated 
measures ANOVA for Left Shoulder (LSPEAK) and Right Shoulder (RSPEAK) was 
conducted to determine if there was a significant difference between the contact pressure 
levels at both shoulder locations.  Results revealed that there were no significant 
differences between the two locations, allowing the author to combine of the shoulder 
data into the peak contact pressure at the shoulder (SHLDRPEAK): (F(2,14) = 2.595, p = 
0.128, ?2 = 0.147, Power = 0.326).  
 
Peak Shoulder and Abdominal Contact Pressures 
 
Following the regression analysis, a 2 (loading condition: SDS and MDS) x 2 
(body location: shoulder and abdominal) x 3(time: pre5, pre55 and post55) multivariate 
repeated measures ANOVA with a p value set apriori at p < 0.05 yielded a significant 
interaction of peak contact pressure between body location and loading condition, 
(F(1,15) = 8.993, p = 0.009, ?2 = 0.375, Power = 0.800).  This indicates that the loading 
conditions altered the magnitude of contact pressures differently across the body 
locations, as well as where the load was distributed during the two loading conditions. 
Descriptive statistics for all dependent variables are presented in Appendix F. 
 92 
In addition to a repeated measures ANOVA across the two loading conditions 
(SDS and MDS) per body placement, two paired t-tests were conducted to investigate 
both loading conditions (SDS and MDS) for SHLDRPEAK and ABSPEAK respectively.  
Each test yielded a significant difference.  For the SHLDRPEAK there was a significant 
difference between the SDS (11.59 kPa) and MDS (8.89 kPa) conditions (p = 0.021), 
with the MDS condition yielding lower pressures.  For the ABSPEAK there was a 
significant difference between the SDS (4.77 kPa) and MDS (5.48 kPa) conditions (p = 
0.009), with the MDS condition yielding higher pressures than the SDS.  These results 
indicated that placing the SDS load onto the participants' body resulted in the largest 
contact pressures at the shoulders but less pressure at the abdominal region due to the 
lumbar belt being implemented to redistribute the load to the abdominal plate, pelvis, and 
belt (Figure 43). 
 Figure 43.  Effects of loading conditions on SHLDR
PEAK and ABSPEAK contact pressure.          
*p = .021. +p = .009. ?p = .001. ?p = .033. 
 
 
0
2
4
6
8
10
12
14
SDS MDS
Co
ntac
t Pr
essur
e (
kPa)
 
Loading Conditions 
Effects of Loading Conditions on 
SHLDRPEAK and ABSPEAK Contact Pressure 
SHLDR
ABS
* 
+ 
* 
+ ? 
? 
? 
? 
 93 
CHAPTER V 
DISCUSSION 
 
 The purpose of this investigation was three-fold:  1) to determine the influence of 
a standard marching snare drum system (SDS) on trunk and lower extremity kinematics 
while standing and marching, 2) to determine the contact pressures and muscular demand 
placed on the erector spinae of individuals wearing a SDS, and 3) to investigate the 
influence of a lumbar belt (SDS-modified (MDS)) on trunk and lower extremity 
kinematics, muscular activity, and contact pressures.  The following chapter has been 
divided into three primary sections: Section 1: The effect of the SDS and MDS systems 
on joint kinematics, Section 2: The effect of the SDS and MDS systems on ES muscle 
activity, and Section 3: Difference in contact pressures between the SDS and MDS 
loading conditions, at the shoulder and abdominal regions.  
Section 4 includes final conclusions and recommendations for future research. 
 
 
 
 
 
 
 
 94 
Section 1: SDS & MDS Conditions on Joint Kinematics 
 
 This section will discuss the results regarding the demands of wearing a SDS and 
MDS on the gait kinematics of marching drumline individuals, compared to an unloaded 
condition.  The gait kinematics of interest are the average angular kinematics of the trunk 
(TAAVG), hip (HAAVG), and ankle (AAAVG).  It should be noted that knee angle was also 
investigated, but no significant differences were found for that particular joint, therefore 
will not be included in this discussion.  The angular kinematics were investigated while 
the participants stood still and played the drum (stationary) as well as when they marched 
forward and played (dynamic).  Table 4 illustrates the results for each joint and 
movement condition across all the two loading conditions. 
 
Table 4 
Effect of Loading Conditions on Joint Kinematics 
 Note.      = Angle larger than U condition, (increased trunk flexion, increased ankle 
plantarflexion, and decreased hip flexion).     = Angle smaller than U condition 
(decreased trunk flexion, and increased ankle dorsiflexion).  * or + = Significant 
difference from U condition 
 
 Previous literature has indicated that adaptations to an external load placed on the 
front of the body include trunk extension which shifts the COGsys closer to an unloaded 
position (Devroey et al., 2007; Ling et al., 2004), and increased hip and knee flexion 
Kinematic Variable SDS MDS 
TA AVG  Stationary 
HA AVG  Stationary 
AA AVG  Stationary 
TA AVG  Dynamic 
HA AVG  Dynamic 
AA AVG  Dynamic 
  * 
 
 
 
 
* + 
+ * 
+ * 
 95 
which improves shock absorption and lowers the COGsys (Anderson et al., 2007; Cham & 
Redfern, 2004; Holbein & Redfern, 1997; Kinoshita, 1985).  However, there are specific 
demands/constraints that are involved with participating in a marching band as a snare 
drummer that do not allow for most of these previously noted adaptations.  The 
constraints include but are not limited to: restricted arm position, standardized step size, 
drum position, and body posture.  As a result, individuals who wear a snare drum and 
harness will experience an anterior shift of the center of gravity of the person (COGperson), 
but will be unable to take advantage of the adaptations of trunk extension, natural arm 
swing, changes in stride length and frequency, nor changes in joint angles in an attempt 
to alleviate the stress being placed on the body.  In observing a marching band practice or 
performance, drumline members seem to be maintaining an upright posture, however the 
trunk measure findings during both load conditions indicate that the drum caused the 
trunk to be in a significantly flexed position (See Figures 31 and 32).  According to 
Keyserling (1986), neutral trunk position occurs at less than 20? from the vertical.  In the 
present study, during both the stationary and dynamic movement conditions, the 
participants' unloaded trunk angles were within those parameters (18.70? and 15.79? of 
flexion respectively).  However, during both the SDS and MDS conditions the TAAVG 
was larger than 20?.  This finding indicates that the external load was great enough to 
alter the trunk kinematics outside what has been considered neutral posture.  This finding 
was surprising as it was anticipated that the anterior load would cause the performer to 
extend the trunk in order to bring the load closer to the base of support.  In addition, it 
was thought that the addition of the lumbar support would attenuate the trunk extension.  
At this point it is unclear whether the trunk flexion is perceived as more aesthetically 
 96 
(determined by the director) pleasing or is an attempt by the performer to utilize the 
upper posterior portion of the harness support.   
There were no significant changes in the hip during the dynamic movement 
condition; however, the trends did show that the MDS loading condition had the largest 
absolute hip flexion angle.  Therefore it can only be speculated that the MDS condition 
allowed for increased ROM (flexion and extension at the hip) and, thus the larger HAAVG.  
The increased hip flexion during the dynamic movement could also be an indication of 
the hip/upper leg being called upon to act as a shock absorber during the load carriage 
conditions.  It has been stated in previous literature that the lower extremities will act as 
shock absorbers during load carriage (Kinoshita, 1985), and normally one would expect 
that the knee would be the most likely joint for this role.  However, in the present study 
there were no significant changes in knee angle while stationary, throughout one stride 
(dynamic), nor during the loading phase (foot flat).  This leads the author to conclude that 
the role of shock absorber for the body is unknown.  It is reasonable to conclude that the 
lack of changes noted at the knee is the result of the restrictive nature of the marching 
gait which requires the participants to have very little knee flexion when making contact 
with the ground and to remain as upright as possible (i.e. no opportunity to lower the 
COGsys closer to the ground) throughout the stance phase.   
There was however, a significant difference in HAAVG between the U and SDS 
conditions during the stationary condition.  The SDS load resulted in a significantly 
larger HAAVG which is indicative of a thigh position closer to vertical and in this 
movement condition indicates a more upright posture.  This does not indicate that 
wearing the SDS causes better posture, but instead suggests to the author that the anterior 
 97 
shift in the COGperson due to the SDS system resulted in the participants leaning back, 
thus increasing the angle of the thigh to parallel.  Without a significant change in the knee 
angle, future research is needed to determine if other factors such as pelvic girdle 
movement could be influencing the hip angle during stationary marching tasks. 
 The effect of loading an individual with either the SDS or MDS systems resulted 
in significant changes in AAAVG, but only during the stationary condition.  It was initially 
hypothesized that the amount of plantarflexion would increase during the stationary 
condition when loaded because of the shift in COGsys forward, resulting in the 
participants attempting to shift the COGsys back and doing so by pivoting about the ankle.   
 It was believed that while marching, the participants would also have an increase 
in dorsiflexion.  The results however, did not find a significant effect of load condition 
for AAAVG, but instead found that all three loading conditions had AAAVG close to the 
fundamental 90? angle.  This indicates that regardless of loading condition, when asked 
to perform the roll-step marching gait, that experienced marchers will automatically place 
the ankle at the required 90?. 
 
Average Knee Angle During Foot Flat (FF) 
 
In addition to investigating the effects of the SDS and MDS loading conditions on 
the joint kinematics of the body during a stride, the author also investigated the average 
joint kinematics of the knee during the FF phase of the stride length (KAAFF).  Results for 
this portion of the study did not yield a significant effect for loading condition (U, SDS, 
and MDS) nor time (pre, post).  This indicates that the loading conditions do not alter the 
joint kinematics of the knee differently during the foot flat phase of the gait cycle.  
 98 
Although not stated in the primary or secondary hypotheses, based on the current 
literature, it was expected that the SDS condition would cause an increase in knee flexion 
during the loading phase of the gait cycle.  This posture would serve as a protective 
mechanism by the participants by lowering the COGperson for increased balance and 
control (Holt et al., 2003; Kinoshita, 1985).  The lack of a finding supports the notion that 
the marching gait constraints supersede the load carriage demands. 
 
Average Ankle Angle During Foot Flat (FF) 
 
In addition to investigating the effects of the SDS and MDS loading conditions on 
the joint kinematics of the body during a stride, the author also investigated the average 
joint kinematics of the ankle during the FF phase of the gait cycle (AAAFF).  Results for 
this portion of the study were similar to the KAAFF data, where no significant effect for 
loading condition (U, SDS, and MDS) or time (pre, post) was found.  This indicates that 
the ankle kinematics were the same for all three loading conditions during the foot flat 
phase of the gait cycle.  The reason for investigating this factor stemmed from the 
relationship between the lower leg and the foot, where it was believed that the SDS 
condition would yield more knee flexion and smaller ankle angle, indicating that the knee 
and ankle structures were acting as shock absorbers during the gait cycle.  However, the 
constraints of the marching gait prevented such adaptation. 
 
 
 
 
 
 99 
Section 2: SDS & MDS Condition on ES Muscle Activity 
In addition to significantly affecting the joint kinematics of marching drumline 
members, load carriage tasks also affect the muscular activity of muscles involved in 
postural stability (i.e. ES).  Therefore, the demands of wearing a SDS and MDS on the 
muscular activity of the ES while standing and playing in place (stationary) and marching 
forward while playing (dynamic) were also investigated.  The average and peak EMG 
levels were compared across all three loading conditions.  The previous section discussed 
the influence of load, movement, and time on the joint kinematics of the body, and it is 
those changes in joint kinematics (whether due to the testing variables or the restrictive 
nature of marching) that influence the demands on the muscles.  Anderson et al. (2007) 
investigated the influence of carrying an anterior load on the body on the muscle activity 
of the ES, while participants stood still and walked.  While holding the various loads 
anterior to the trunk, the researchers documented increased trunk extension and 
subsequent increases in ES muscle activity.  This supported the development of the 
hypothesis that placing a load on the anterior portion of the body would result in 
increased muscle activity of the ES in order to maintain the erect body posture required 
by the demands of this type of performance.  Table 4 illustrates all of the significant 
findings for the ES during each movement condition across all three loading conditions. 
 
 
 
 
 
 
 
 
 
 
 100 
Table 5 
Effect of Loading Conditions on sEMG of the ES 
 
Note.      = Muscle activity larger than U condition. *or + = Significant difference from U 
condition 
 
It can be concluded from the present project that load carriage, particularly either 
the SDS or MDS systems, affect levels of EMGAVG and EMGPEAK ES activity in 
collegiate male and female marching drumline members.  This supports the finding of 
Bobet and Norman (1984) and Hamilton et al. (2008), which indicated that deviations 
from a normal, unloaded, anatomically neutral posture can result in a change in muscular 
demand.  An additional demand that is placed upon the postural musculature is the 
inability to utilize arm movement for balance control, due to performance constraints.  
Natural arm movements modulate the excursions of the COGsys (Birrell & Haslam, 
2008), and therefore, when those movements are limited or restricted, the result can be 
greater excursions of the COGsys within the BOS and therefore decreased stability.  In 
order to compensate for the instability, stabilizing muscle activity (i.e. postural muscles) 
will increase, resulting in an earlier onset of fatigue or "fatigue-like effects" (Cook & 
Neumann, 1987).  
Al-Khabbaz et al. (2008) determined that placing a load on the posterior part of 
the body (10, 15, and 20% BM) would result in an increase in muscle activity (ranging 
from 3-6% MVIC) of the rectus abdominus (RA) in an attempt to counteract the eccentric 
sEMG-ES Activity SDS MDS 
EMG AVG   Pre  
EMG AVG   Post 
EMG PEAK  Stationary 
EMG PEAK  Dynamic 
 
   
  
  
  
* 
* 
+ 
* 
* + 
+ 
+ 
 101 
trunk extension moment.  Motmans et al. (2006) also found that during a stationary 
posterior load carriage condition, the muscle activity of the ES decreased while the RA 
increased as a result of the extension moment.  Interesting to note, even though Al-
Khabbaz et al. (2008) determined that the RA activity had increased to compensate for 
the posterior load and backward inclination, the trunk was still in an extended position.   
Similar to the findings of Al-Khabbaz et al. (2008), Foti, Davids & Bagley (2000) and 
Motmans et al. (2006), the present study significant increases in muscle activity in the 
presence of a change in trunk position.  For this particular study, it was a significant 
change in ES activity for the change in trunk flexed position.    
For the current study, EMGAVG was analyzed before (results ranging from 1.7-
3.6% MVIC) and after a one-minute exercise (results ranging from 1.72-4.01% MVIC), 
similar to the findings of Al-Khabbaz and colleagues (2008) of 2.96-5.37% MVIC for 0-
15% BM loads.  The researchers believed that this increase in muscle activity was an 
attempt to counteract the eccentric trunk flexion moment that was occurring due to the 
anterior loading conditions.    Similar results were found for the EMGPEAK where during 
the stationary trials (results ranging from 3.05-5.61% MVIC) and the dynamic trials 
(results ranging from 5.62-10.25% MVIC), significant differences were found between 
the U and both SDS and MDS conditions.  Although these levels of muscle activity were 
not the same amount of activity to the findings of Anderson et al. (2007) (results ranging 
from 30s-40s% MVIC), the data do indicate that there is a significant difference in 
muscle activity when there has been a change in postural demand due to an external load.  
Overall, the effect of placing an external load on the body results in postural demands as 
the COGsys shifts anteriorly.  Changes in posture result in changes in muscle activity, 
 102 
specifically the ES attempting to shift the COGsys more within the BOS by extending the 
trunk.   During both the stationary and dynamic trials, before and after the marching 
exercise, there was a significant difference in muscle activity of the ES for both the SDS 
and MDS conditions from an unloaded position.  Therefore, it is reasonable to conclude 
that the participants' postures was influences by being loaded with a snare drum and 
carrier, and that before the marching exercise, the participants were not utilizing the load 
distribution characteristics of the lumbar belt.   After the marching exercise, there were 
still significant differences for the stationary and dynamic tasks for both the SDS and 
MDS conditions and no significant difference between the two loads.  However, there 
was a trend of decreasing muscular demand during the MDS to indicate that the 
participants might be making subtle changes in posture, therefore allowing the 
participants to utilize the load distribution characteristics of the lumbar belt.         
 
Section 3: SDS & MDS Condition on Contact Pressure 
The purpose of this portion of the current study is to present the findings and 
implications of peak contact pressures at the shoulders and abdominal region.  These data 
were collected while wearing a SDS and MDS in a stationary position for one minute 
while playing, prior to and after a one minute marching exercise.  The pressure data were 
taken at three times: Pre5-during the initial 5 seconds of the stationary/playing loading 
one minute period prior to the one minute marching exercise, Pre55-during the last 5 
seconds of the stationary/playing loading one minute period prior to the one minute 
marching exercise, and Post55-during the last 5 seconds of the stationary/playing one 
minute period after the one minute exercise.  The data of interest included the peak 
 103 
contact pressure values from both loading conditions (SDS and MDS), at both the 
shoulders (SHLDRPEAK) and abdomen (ABSPEAK).  Table 5 illustrates the significant 
findings for the contact pressures for each loading condition, across both body segments.   
Table 6 
Effect of Loading Conditions on Contact Pressure 
 
Note.      = Highest contact pressure. * = Significant difference between SDS and MDS.     
 
  Skin contact pressure is considered one of the primary limiting factors in the 
ability to perform a load carriage task (Holewijn, 1990), particularly if the load is not 
removed or reduced after a period of time.  Doan (1998) determined that levels as low as 
14 kPa can be a limiting factor, even though the participants in that study were not 
reporting any discomfort until 18-20 kPa.  Lower pressures however do not necessarily 
mean that damage is not occurring in the tissues or nervous system (Goodson, 1981), nor 
should individuals avoid acknowledging the discomfort.  The current study collected 
SHLDRPEAK levels between the SDS (11.59 kPa) and MDS (8.89 kPa) conditions, and 
ABSPEAK levels between the SDS (4.77 kPa) and MDS (5.48 kPa) conditions, which are 
lower than the limiting value of 14 kPa determined by Doan (1998) and the performance 
criterion of 20 kPa indicated by Stevenson et al., 2004.   
  The present study also found a statistically significant decrease in pressure 
between the SDS and MDS configurations, similar to Holewijn (1990) and Mackie et al. 
(2005) who noted that implementing a lumbar belt significantly changes the distribution 
and amount of pressure being applied to the body during a load carriage condition.  
Contact Pressure SDS MDS 
SHLDR PEAK 
ABS PEAK 
* 
* 
 104 
Adding a lumbar belt to the SDS system not only reduced the amount of pressure at the 
shoulders, but also redistributed the forces to the abdominal region and pelvis which are 
areas better suited for load carriage (Scribano, 1970) (Figure 43). Anecdotally, 
participants remarked about the discomfort of the SDS system, but not the MDS system.  
Specific comments from the participants in a post data collection interview included: 
"Alleviates the pressure between the shoulder blades," "Takes away the pressure off the 
lower back right away," "Belt takes the weight off shoulders and off of you," "Doesn't 
push into stomach, just sits against it," and "Takes the pressure/weight off shoulders and 
puts on back and abs where the belt which feels good."  
 
Overall Impact 
Loading an individual with an anterior load, even a participant with experience in 
that particular loading condition, causes significant changes in joint kinematics, muscle 
activity, and contact pressure and the inclusion of a lumbar belt does not diminish the 
changes in kinematics and muscle activity.  However, significant differences were noted 
in contact pressure findings.  Implementing a lumbar belt did redistribute the contact 
pressures, changed the location of the drum with respect to the drummer, and most 
importantly did not harm the participants but instead made wearing a drum "feel better."   
It was also discussed that the inclusion of the lumbar belt would decrease 
independent motion of the drum.  One of the objectives of the lumbar belt was to make 
the drum and harness more connected to the performer and therefore decreases the drum 
from behaving as an independent inertial body.  As an independent inertial body, the 
drum becomes like a person in a car without a seatbelt, the lumbar belt was thought to be 
 105 
the ?seatbelt?.  While this objective was difficult to measure quantitatively, the 
participants provided the following comments:  "Cohesive," "Keeps the drum from 
bouncing around," "More aware of where the load is and how the drum moves, ""Now I 
have to focus on really rolling my feet properly so that the drum doesn't bounce," and 
"The belt brought the drum closer to the body, made it easier to move, and easier to 
perform maneuvers." 
 
Section 4: Summary & Future Research 
 
 This final section will discuss the conclusions drawn from the current study and 
directions for future research.  The goal of this study was investigate the loading effects 
of a traditional marching snare drum and carrier system on the kinematics, muscle 
activity, and contact pressures of the body.  In addition to the baseline study, the author 
also investigated implementation strategies demonstrated by previous literature to 
alleviate the stress of load carriage on the body.  The end result was the creation and 
construction of a lumbar belt specifically designed for marching drumline carriers, which 
was tested to determine changes in kinematics, muscle activity and contact pressures on 
the body.   
When an individual wears a standard marching snare drum system while standing 
still, there are several significant changes: increased trunk flexion; decreased hip flexion; 
increased ankle plantarflexion (most likely as a result of the ankle acting as the axis of 
rotation for the large lever arm of the whole body and attempting to shift the COGperson 
posteriorly); increased ES EMGAVG and EMGPEAK as a result of the changes in trunk 
 106 
angle; and increased contact pressures at the shoulders indicating that the load is still 
predominately carried at the shoulders. 
When an individual wears a standard marching snare drum system while 
marching, the results showed significant: increased trunk flexion, and increased ES 
EMGAVG and EMGPEAK as a result of the changes in trunk angle.  
 When implementing a lumbar belt there were similar significant findings at the 
joints and muscle activity level from the unloaded condition only, suggesting that the 
implementation of the lumbar belt did not alter the marching kinematics of the performer.  
However, a significant difference was noted between the two systems for contact 
pressure.  It was determined that the MDS system caused lower contact pressure values at 
the shoulders and higher pressures at the abdominal region, thus indicating a 
redistribution of pressure and load to the belt, hips, abdomen, and pelvic girdle and off of 
the shoulders.  Therefore it can be concluded that implementing a lumbar belt to a SDS 
system ameliorates skin contact pressure and discomfort at the shoulders, potentially 
decreasing the occurrence of an injury, and anecdotally providing relief to the 
participants.   
 
Future Research 
When considering load carriage, the primary variable of concern is the size of the 
load.  Researchers have found significant changes in gait kinematics, muscle activity, and 
contact pressure with load sizes as low at 10% body mass (BM), and larger and more 
severe changes occurring at 20% BM.  These findings have caused the literature to 
recommend that load carriage be limited to 20% BM (Devroey et al., 2007; Kinoshita, 
 107 
1985; Malhotra & Sen Gupta, 1965; Pascoe & Pascoe, 1999).  The results of this study 
found that the SDS system was on average 11.1% BM (ranging from 5.8-15.1% BM of 
the participants), explaining the significant findings of this study, while not exceeding the 
recommended load carriage limit of 20% for this particular population.  However, the 
current study only investigated the effects of an SDS on collegiate marching band 
members.  Yet this SDS is utilized by thousands of high school and junior high marching 
band participants and it is quite possible that this load could reach or exceed the 
recommended limit of 20% BM.  Therefore, future research should be conducted on these 
younger populations to determine the effect of the SDS and MDS on these smaller 
performers.   
 Significant increases in muscle activity were found when Hong et al. (2008) 
investigated the effects of not only load carriage mass but load carriage duration.  It was 
determined that the larger the mass the sooner it took for the muscle activity to reach a 
significantly different and fatigued level (i.e. 5 minutes at 20% BM, and 15 minutes at 
15% BM respectively) .  The EMGAVG data from the present study demonstrated 
increases in muscle activity levels from pre to post a one minute marching exercise.   
Therefore variable of time (i.e. duration of wearing the SDS) will need to be addressed in 
future studies to determine duration effects of load carriage, adaptations or any learning 
effects, with the potential to identify efficient practice protocols.   
In addition to these areas of future research, investigations should also be 
expanded to include 1) the other anterior load carriage scenarios in marching bands (i.e. 
quad/tenor drums, base drums, and cymbals), 2) vertical/direct loading to the spine in 
marching band (i.e. tubas), and 3) different populations of marching band participants 
 108 
(i.e. junior high, high school, and drum corps).  As well as the long term effects (i.e. 
longitudinal studies) on participants in marching drumlines, based on anecdotal evidence 
of recurring injuries, pain, and discomfort attributed to being a marching drumline 
participant.  These individuals, after all, are as passionate about this activity as an athlete 
is about their sport.  However, less attention has been paid to the side effects of 
participation in marching activities compared to traditional athletic endeavors.   
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 109 
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biomechanics of sit-to-stand motion in healthy children. Human Movement 
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Sharpe, S. R., Holt, K. G., Saltzman, E., & Wagenaar, R. C. (2008). Effects of a hip 
belt on transverse plane trunk coordination and stability during load carriage. 
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Shasmin, H. N., Abu Osman, N. A., Razali, R., Usman, J., & Wan Abas, W. A. B. 
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116 
 
Appendix A 
 
Preliminary Medical Questionnaire 
 
Please read each question carefully and answer honestly.  If you do not 
understand the question, please ask the investigator for clarification.  Check the 
appropriate answer. 
    
Participant Number: ___________ 
 
   YES     NO 
 
                          1) Are you under the age of 19? 
 
                          2) Have you ever been told you have an inner ear disorder? 
 
                          3) Have you ever had lower extremity surgery? 
 
                          4) Do you presently have any lower extremity disorders? 
 
_____  ____   5) Has your doctor ever told you that you have a muscle, bone or joint  
problem such as arthritis that has been aggravated by exercise, 
or might be made worse by exercise? 
                              
                          6) Have you ever felt faint, dizzy or passed out during or after  
exercise? 
 
                          7) Have you ever felt pain, pressure, heaviness or tightness in the  
chest, neck shoulders or jaws as a result of exercise? 
 
                          8) Do you have any reason to believe that your participation in this  
investigative effort may put your health or well-being at risk? 
 
                          9) Have you had an injury or surgery to either of your lower  
extremities? 
 
                          10) Are you able to walk and stand on one leg without the use of  
aids?  
 
                          11) Are you currently taking any medication that you think might  
influence your ability to participate in this study?  
117 
 
Appendix B 
 
Institutionally Approved Informed Consent Document 
 
INFORMED CONSENT  
for a Research Study entitled 
 
The Influence of a Marching Snare Drum Harness on Joint Kinematics, 
Electromyography and Contact Pressure 
 
You are invited to participate in a research study to look at the load 
carriage demands placed on the collegiate marching band athlete.  This study is being 
conducted Andrea Sumner and the supervision of Dr. Wendi Weimar, Associate 
Professor of Biomechanics, in the Auburn University Department of Kinesiology.   
You were selected as a potential participant because you are between the ages of 19 
and 25 years of age, and your health condition and mobility might, through pre-
screening health questionnaire to follow, permit you to perform these tests safely and 
successfully.  The subject population is that of convenience and who represent the 
current collegiate marching band athlete.  The results of this study will be used to 
investigate the biomechanical effects on the trunk and lower extremities associated 
with wearing a snare drum carrier and drum, and how the implementation of a lumbar 
belt to the snare drum carrier influences these effects. 
What will be involved if you participate?  If you decided to participate in 
this experiment, your testing session will last approximately 45 minutes, which will 
include time to fill out forms.  After meeting the criteria for participation, both your 
mass and height will be measured for normalizing purposes.  You will then be asked 
to change into spandex shorts and spandex tank tops to minimize clothing artifact.  In 
order to identify the muscle group, the researchers will ask you to perform a trunk 
extension (bending backwards) movement.  After the joint movement, the skin area 
will be cleaned with a 70% alcohol solution and abraded to reduce the electrical 
impedance of the skin.  A ground electrode will be placed on the bony aspect of the 
olecranon process (elbow) and two additional electrodes will be placed on the 
muscles at your lumbar 4-5 interspace.  Next, MVIC (maximum voluntary isometric 
contraction-used for normalization) will be collected for the group. 
Following the sEMG (electromyography-muscle activity) electrode 
placement, three pressure sensors will be laid over the designated landmarks.  Two
118 
 
pressure sensors (#9811E) will be placed directly under the location where the drum 
carrier makes contact with the body.  The sensors will lie over the superior (top) aspect of 
the shoulders, angled towards the midline, and cross over the clavicle.  The third sensor 
(#9833E) sensor will be placed across the lower abdominal/pelvic region, directly under 
the location where the drum carrier makes contact with the front of the body with the 
center of the sensor at the midline of the body.               
 
Finally, retroreflective markers will be affixed to the spandex clothing and to the 
skin with double-sided tape and placed according to the requirements of the Vicon? 
Plug-In-Gait? model.  A static capture will be performed while you stand in the center of 
the capture volume in an anatomical position with shoulders out to the side.  Once the 
static capture is performed, you will begin the data collection.  The procedures for this 
study are the same throughout the study, with the only changes occurring for the different 
loading conditions.  The loading conditions consist of unloaded, wearing your snare drum 
and harness without any modifications, and wearing your snare drum and harness with 
the addition of a lumbar belt.  The procedures for this study for each loading condition 
are as follows: 1) you will be asked to stand for 1 minute while sEMG, pressure data, and 
joint kinematics are collected, 2) you will then march (glide step and crab step gait) 
through capture volume while sEMG, pressure, and joint/gait kinematic data is collected, 
3) you will then march (glide and crab step gait) for 10 minutes in the coliseum 
gymnasium, 4) you will then reenter the capture volume and stand in a stationary position 
for 1 minute while EMG, pressure data, and joint kinematics will be collected once again, 
and 5) you will then march (glide step and crab step gait) through capture volume while 
sEMG, pressure, and joint/gait kinematic data is collected. 
During the static trials, your arms will be in the playing position and your feet will 
be at a 45? angle.  During the unloaded condition, the arms will be in the playing position 
as well in order for control for arm placement throughout the conditions, as well as to 
mimic the requirements that occur during practice of ?air playing.? There will be a period 
of 5 minutes of rest in between each loading condition, and more will be afforded to you 
if needed.  
   Are there any risks of discomforts?  There is a slight risk associated with 
walking, marching, and carrying an instrument.  However, since these are usual activities 
for you as a collegiate marching band member, these risks are minimal.  Last, a Certified 
First Aid/First Responder will be present at all testing sessions.  If medical treatment is 
necessary due to an injury, you are responsible for any and all medical costs. 
Are there any benefits to yourself or others?  As a participant, it is hoped that 
you will have a better understanding as to how load carriage effects your gait (if it does), 
and how the demands placed on your body is affected by the current harness system you 
have and one that had been modified.  Hopefully this information will provide the 
researchers with ideas that can help you as an athlete to be more efficient on the field and 
less susceptible to injuries. We/I cannot promise you will receive any of the benefits 
described. 
Will you receive compensation for participating?  To thank you for your time, 
you may request the data.  There will be no compensation for participating in this project.   
If you change your mind about participating, you can withdraw at any time 
during the study.  Your participation is completely voluntary.  If you choose to withdraw, 
119 
 
your data can be withdrawn as long as it is identifiable.   Your decision about whether or 
not to participate or to stop participating will not jeopardize your future relations with 
Auburn University, the Department of Kinesiology, or the Sport Biomechanics 
Laboratory. 
Your privacy will be protected.  Any information obtained in connection with 
this study will remain confidential.  Information obtained through your participation may 
be published in a professional journal or presented at a professional meeting. 
If you have questions about this study, please ask them now or contact Andrea 
Sumner at sumneam@auburn.edu, or Dr. Wendi Weimar at weimawh@auburn.edu.  
Furthermore, you may reach any of these three individuals by calling the Sport 
Biomechanics Laboratory at 844-1468.  A copy of this document will be given to you to 
keep. 
If you have questions about your rights as a research participant, you may 
contact the Auburn University Office of Human Subjects Research or the Institutional 
Review Board by phone (334)-844-5966 or e-mail at hsubjec@auburn.edu or 
IRBChair@auburn.edu. 
 
   HAVING READ THE INFORMATION PROVIDED, YOU MUST DECIDE 
WHETHER OR NOT YOU WISH TO PARTICIPATE IN THIS RESEARCH 
STUDY. YOUR SIGNATURE INDICATES YOUR WILLINGNESS TO 
PARTICIPATE. 
 
____________________________                 ______________________________ 
Participant's signature  Date                  Investigator obtaining consent    Date 
____________________________                _______________________________ 
Printed Name                  Printed Name 
120 
 
Appendix C 
 
ANOVA Table for the Effect of Load Carriage on Joint Kinematics 
 
Stationary   Mean SD p ?2 Power 
TAAVG (?) 
      
 
U 18.70 3.39 
   
 
SDS 21.82 4.55 < 0.001 0.499 0.996 
 
MDS 21.94 4.33 
   HAAVG (?) 
      
 
U 73.82 4.06 
   
 
SDS 77.08 4.14 0.020 0.228 0.719 
 
MDS 78.09 6.38 
   AAAVG (?) 
      
 
U 91.35 3.09 
   
 
SDS 93.75 3.24 < 0.001 0.684 1.000 
 
MDS 93.76 3.34 
   
       
       Dynamic 
 
Mean SD p ?2 Power 
TAAVG (?) 
      
 
U 15.79 3.73 
   
 
SDS 22.96 5.43 < 0.001 0.855 1.000 
 
MDS 22.71 5.10 
   HAAVG (?) 
      
 
U 86.48 2.49 
   
 
SDS 88.30 3.16 0.018 0.236 0.739 
 
MDS 88.72 3.49 
   AAAVG (?) 
      
 
U 90.35 3.61 
   
 
SDS 90.51 3.06 0.475 0.048 0.167 
 MDS 89.80 3.29       
 
 
 
121 
 
Appendix D 
 
ANOVA Table for the Effect of Load Carriage on Joint Kinematics 
 
Foot Flat   Mean SD         
KAAFF (?) 
       
 
U Pre 176.45 1.76 
    
 
SDS Pre 175.80 2.80 
    
 
MDS Pre 175.17 2.94 
    
        KAAFF (?) U Post 176.21 2.67 
    
 
SDS Post 175.54 2.62 
    
 
MDS Post 175.47 2.43 
    
        ANOVA for Loading Condition: (F(2,14) = 1.114, p = 0.341, ?2 = 0.069, Power = 
0.227) 
ANOVA for Time: (F(2,14) = 0.125, p = 0.729, ?2 = 0.008, Power = 0.063) 
        Foot Flat   Mean SD         
AAAFF (?) 
       
 
U Pre 87.92 3.81 
    
 
SDS Pre 87.51 2.90 
    
 
MDS Pre 87.53 2.81 
    
        AAAFF (?) U Post 88.48 3.98 
    
 
SDS Post 88.56 4.21 
    
 
MDS Post 87.40 2.26 
    
        ANOVA for Loading Condition: (F(2,14) = 0.467, p = 0.631, ?2 = 0.030, Power = 
0.119) 
ANOVA for Time: (F(2,14) = 0.691, p = 0.419, ?2 = 0.044, Power = 0.122)   
 
 
122 
 
Appendix E 
 
ANOVA Table for the Effect of Load Carriage on Average ES sEMG 
 
Pre   Mean SD p ?2 Power 
EMGAVG (%MVIC) 
      
 
U 1.70 1.11 
   
 
SDS 3.57 1.73 < 0.001 0.674 1.000 
 
MDS 3.64 1.67 
   
       Post 
 
Mean SD p ?2 Power 
EMGAVG (%MVIC) 
      
 
U 1.72 1.03 
   
 
SDS 4.01 1.92 < 0.001 0.710 1.000 
 MDS 3.69 1.65       
 
 
ANOVA Table for the Effect of Load Carriage on Peak ES sEMG 
 
Stationary   Mean SD p ?2 Power 
EMGPEAK (%MVIC) 
      
 
U 3.05 1.17 
   
 
SDS 5.61 2.33 < 0.001 0.665 1.000 
 
MDS 5.30 2.52 
   
       Dynamic 
 
Mean SD p ?2 Power 
EMGPEAK (%MVIC) 
      
 
U 6.62 3.60 
   
 
SDS 9.74 5.37 < 0.001 0.654 1.000 
 MDS 10.25 5.25       
 
 
 
 
 
 
 
 
123 
 
Appendix F 
 
ANOVA Table for the Effect of Load Carriage on Peak Contact Pressure 
 
    Mean SD       
SHLDRPEAK (kPa) 
      
 
SDS 11.59 5.80 
   
 
MDS 8.89 5.16 
   
       ABSPEAK (kPa) SDS 4.77 1.30 
   
 
MDS 5.48 1.47 
   
       ANOVA for Body Location x Loading Condition :  
(F(1,15) = 8.993, p = 0.009, ?2 = 0.375, Power = 0.800) 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
124 
 
  
 
 
 
 
 
  
Appe
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125 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
 
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126 
 
 
 
 
 
 
  
Appe
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127 
 
 
 
 
 
 
 
  
Appe
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128 
 
 
 
 
 
 
 
  
Appe
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129 
 
 
 
 
 
 
 
  
Appe
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130 
 
 
 
 
 
  
Appe
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131 
 
 
 
 
 
  
Appe
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132 
 
 
 
 
 
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