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Applied Sciences: Biodynamics

Gender Differences in Head–Neck Segment Dynamic Stabilization during Head Acceleration

TIERNEY, RYAN T.1; SITLER, MICHAEL R.1; SWANIK, C BUZ1; SWANIK, KATHLEEN A.1; HIGGINS, MICHAEL2; TORG, JOSEPH3

Author Information
Medicine & Science in Sports & Exercise: February 2005 - Volume 37 - Issue 2 - p 272-279
doi: 10.1249/01.MSS.0000152734.47516.AA
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Abstract

Acute and long-term concussion sequelae are produced by acceleration or deceleration of the freely moving head (10,12). Based on athlete exposure epidemiological data in the late 1990s, females in high school and college had a higher incidence of concussions compared with their male counterparts (e.g., soccer, baseball and softball, and basketball) (8,23). The reason for the higher rate of concussions among females in noncontact sports is unknown (18) but may be due to differences in head–neck segment mass, dynamic stabilization (i.e., muscle activity, timing, strength, and stiffness qualities) compared with males (26).

Females have less head–neck segment mass than males (22), which results in a greater risk of deleterious segment angular acceleration (16) and concussion during standardized force application (26). Dynamic joint stabilization is defined as the ability of the myotendon unit to absorb external loads and minimize excessive joint movement (17). Two primary dynamic stabilizers of the head and neck are the sternocleidomastoid (SCM) and trapezius I (29). The timing and amount of activity of these muscles in response or before the application of an external force should reduce the resultant head acceleration (2,13,15,31). Although there have been no reported gender differences in muscle activity amount during head accelerations, previous authors reported faster neck muscle reflex times in females during low-speed rear-end car impacts (4). Insufficient muscle strength could also predispose individuals to concussion because they would not be able to create the internal (muscle) forces necessary to counter the external forces that result in head acceleration (16). Head–neck segment muscle strength gender differences have been reported for college-aged volunteers with males having 30–40% more isometric cervical flexor and extensor strength (9). The differences were attributed to neck girth (muscle mass) and not muscle function.

Contraction of these primary stabilizing muscles and greater girth also increases muscle and joint stiffness (19,20,32). Neck muscle contraction before external head loading increases resistance to movement (24) and should enhance an athlete’s ability to absorb external forces (2,28,32). In a comparison of one female and seven males (18–23 yr old), however, Reid et al. (24) reported that neck muscle resistance to impact was less in the female under various loads and conditions (e.g., muscles relaxed, muscles preactivated) (24). This is consistent with other authors who reported lower stiffness values at other joints in females compared with males (3,21,30).

Although previous studies have identified gender differences in anthropometric, neuromuscular, and strength variables, no research has examined whether these differences make females more susceptible to greater head acceleration. The purpose of this study was to determine whether gender differences exist in kinematic and dynamic stabilization variables responses to an external force applied to the head.

METHODS

Research Design

The study consisted of a three-factor research design with repeated measures. The independent variables were gender (female vs male), force application (known vs unknown), and force direction (forced flexion, trapezius I, vs forced extension, SCM). The dependent variables consisted of anthropometric, kinematic (peak angular acceleration and angular displacement), EMG (muscle activity peak, area, and onset), stiffness, and isometric strength. An external force applicator was used to apply an external force to the head–neck segment (Fig. 1).

FIGURE 1—EMG trace example. This is an example of a typical EMG trace illustrating SCM onset latency during an unknown force application trial for forced extension.
FIGURE 1—EMG trace example. This is an example of a typical EMG trace illustrating SCM onset latency during an unknown force application trial for forced extension.

Participants

Forty physically active males (N = 20, age = 26.3 ± 4.3 yr, height = 177.1 ± 6.1 cm, mass = 84.5 ± 11.8 kg) and females (N = 20, age = 24.2 ± 4.1 yr, height = 165.3 ± 5.3 cm, mass = 59.0 ± 5.1 kg) participated in the study. Physically active was defined as performing a minimum of 30 min of exercise five or more times a week. One female did not participate in the force application trials due to a scheduling conflict. Also, EMG and stiffness data for three subjects (two males and one female) were omitted because of improper set-up with the load cell during one day of testing. Potential participants were excluded from the study if they had a history of neurological disorder (e.g., seizures), prior cervical spine or head injury (i.e., concussion), or participated in a neck-strengthening program in the 6 months before data collection. Institutional review board approval as well as participant written informed consent and consent to videotape were obtained before data collection.

Instrumentation

Head–neck segment anthropometric assessments.

Participant height, weight, head–neck segment length, and neck girth were assessed. Participant weight was measured in pounds and converted to body mass (kg). Body mass was multiplied by the gender specific head–neck segment to total body mass percentage (male = 8.26% and female = 8.20%) to determine head–neck segment mass (22). Head–neck segment length was measured with a metric tape measure from the seventh cervical vertebrae spinous process to the top of the head with the participant looking at an object at eye level. Neck girth was measured with a metric tape measure just above the thyroid cartilage. The investigator’s intratester reliabilities for the anthropometric measurements were intraclass correlation coefficients (ICC; 2,1) of 0.99 (height), 0.99 (weight), 0.98 (head–neck segment length), and 0.99 (neck girth).

Head–neck segment kinematic assessment.

The PEAK Motus Motion Analysis System (Peak Performance Technologies, Inc., Englewood, CO) was used to collect two-dimensional kinematic data. Video was collected at 60 Hz, and reflective markers were used to aid in the digitizing process. The head–neck and torso segments (11) were used to determine head–neck flexion and extension peak angular displacements (°) and accelerations (°·s−2). Raw video data were autodigitized, filtered (fourth-order, zero-lag Butterworth filter with a 6-Hz cutoff), and analyzed using the PEAK Motus software, version 6.1 (Peak Performance Technology’s, Inc., Englewood, CO). Intratester measurement reliability of this instrument has been reported to be an ICC of 0.98 (27).

Head–neck segment EMG assessment.

The Noraxon Telemyo System (Noraxon U.S., Scottsdale, AZ) was used to assess the EMG activity of the SCM and trapezius I muscles. They were chosen because of their importance as superficial muscles that help to control head–neck flexion and extension and utilization in previous head and neck research (2,32). The skin over the right SCM and trapezius I muscles was shaved, lightly abraded, and cleaned with 70% alcohol. Ten-millimeter–diameter self-adhesive silver/silver-chloride bipolar surface electrodes (Multi Bio Sensors Inc., El Paso, TX) were placed on the skin 10 mm apart and parallel to the fiber orientation of the underlying muscle. The resistance between the paired electrodes was less than 2 kΩ and verified with a standard digital multimeter (model 982017, Sears, Roebuck & Company, Hoffman Estates, IL). Placement of the electrodes was identified by palpating the midlength of the muscle’s contractile component during an isometric contraction. A reference electrode was positioned on the skin over the right clavicle. Signals from the muscle leads were passed to a battery operated eight-channel FM transmitter worn by the participant. The signal was amplified (gain 1000) with a single-ended amplifier (impedance > 10 MΩ) and filtered with a fourth-order Butterworth filter (10–500 Hz) and common mode rejection ratio of 130 dB at direct current (minimum 85 dB across entire frequency of 10–500 Hz). An antenna receiver (Antennex, Inc., Glendale, IL) with a sixth order filter (gain 2, total gain 2000) further amplified the signal. The analog signal was converted to a digital signal by an analog-to-digital converter card (Keithley KPCMCIA 12A1-C, Keithley Instruments, Inc., Cleveland, OH) and was stored in the MyoResearch Software, version 2.02 (Noraxon U.S.). The raw digital signal (for MVC and trials) was sampled at a rate of 960 Hz, rectified, and smoothed using a root mean square algorithm over a 20-ms moving window. All analyses were performed on processed EMG data during a 250-ms time period after force application. This time period was chosen because pilot data of reactive muscle activity revealed the greatest activity within 250 ms after force application.

Data collected with the EMG equipment were used to determine peak muscle activity (%), muscle activity area (%·ms), and muscle onset latency (ms). Peak muscle activity was defined as the highest amplitude of the smoothed EMG data during one trial. Muscle activity area was defined as the product of the sum of the amplitudes of EMG activity and the total time of the trial (250 ms). The data for the two muscle activities were normalized to a peak value obtained during a MVC for each muscle. Subjects maintained a neutral head position (i.e., head up and eyes facing forward, looking straight ahead) during the maximum voluntary contraction assessment. Muscle onset latency was defined as the time between force application and the first upswing of myoelectric activity from baseline (5) and measured only during the unknown force application trials (Fig. 1). The primary investigator’s intratester measurement reliabilities for the EMG dependent variables were ICC of 0.92 (peak activity), 0.87 (muscle activity area), and 0.72 (onset latency).

Head–neck segment stiffness assessment.

A tension force load cell (model ELFS-T3, Entran Devices, Inc., Fairfield, NJ) and the Peak Motus Motion Analysis System were used to assess head–neck segment stiffness. Tension force was assessed during the entire trial using the load cell that was inserted in-line with the external force applicator’s pulley cord and headgear (Fig. 2). Head–neck segment angular displacement was assessed using the Peak Motus Motion Analysis System. Tension force and head–neck segment displacement data were synchronized during each trial using the time of force application as the event that initiated data collection. The peak force during each trial was used to mark the end of the data collection. This enabled a force-angular displacement line to be created. Head–neck segment stiffness was determined from the slope of the line (3) that was defined as the change in force (lb) over the change in angular position (°) (19) (Fig. 3). The reliability for the amount of force applied using the tension load cell was an ICC of 0.98 (force application).

FIGURE 2—External force applicator. The external force applicator was designed and constructed by the investigator. The applicator consists of a metal outer frame, headgear, two cords with plastic stoppers, and two pulleys (only forced extension illustrated). Cords from the front or back of the headgear (Strength Systems Inc., Jefferson, LA) wrap around the pulleys and connect to plastic stoppers (landing surfaces) at the end of the cord opposite the headgear. The plastic stopper enables a 1-kg mass to be dropped from a predetermined height of 15 cm, creating a load of approximately 11 lb (50 N). The load is verified during testing using a tension force load cell (model ELFS-T3, Entran Devices, Inc., Fairfield, NJ). The pulleys in front and back of the seated participant are used to cause head–neck segment flexion and extension, respectively. The heights of the pulleys are modified so that each force is applied at 90° to the head–neck segment of the participant during testing. This is verified by visual inspection.
FIGURE 2—External force applicator. The external force applicator was designed and constructed by the investigator. The applicator consists of a metal outer frame, headgear, two cords with plastic stoppers, and two pulleys (only forced extension illustrated). Cords from the front or back of the headgear (Strength Systems Inc., Jefferson, LA) wrap around the pulleys and connect to plastic stoppers (landing surfaces) at the end of the cord opposite the headgear. The plastic stopper enables a 1-kg mass to be dropped from a predetermined height of 15 cm, creating a load of approximately 11 lb (50 N). The load is verified during testing using a tension force load cell (model ELFS-T3, Entran Devices, Inc., Fairfield, NJ). The pulleys in front and back of the seated participant are used to cause head–neck segment flexion and extension, respectively. The heights of the pulleys are modified so that each force is applied at 90° to the head–neck segment of the participant during testing. This is verified by visual inspection.
FIGURE 3—Stiffness calculation example. This is an example of a force-displacement line created from a trial with three force (Y1, Y2, and Y3) and three displacement (X1, X2, and X3) data points. These points create two lines and stiffness was determined from the average slope of the lines
FIGURE 3—Stiffness calculation example. This is an example of a force-displacement line created from a trial with three force (Y1, Y2, and Y3) and three displacement (X1, X2, and X3) data points. These points create two lines and stiffness was determined from the average slope of the lines:
(3).

Head–neck segment isometric strength assessment.

The Microfet Hand-Held Dynamometer (Hoggan Health Industries, Inc., West Draper, UT) was used to assess head–neck segment isometric flexor and extensor muscle strength. Flexor strength was assessed with the dynamometer placed in the center of the participant’s forehead. Extensor strength was assessed with the dynamometer placed just above the participant’s external occipital protuberance. The participant applied maximum force against the dynamometer for 3 s during each trial and rested for 30 s between trials. The primary investigator’s intratester measurement reliability for this instrument was an ICC of 0.96.

Procedures

Potential participants met with the primary investigator and the purpose and procedures of the study were explained. Before testing, participants performed a neck warm-up consisting of neck rotations (15 s clockwise and 15 s counter clockwise) and neck stretching (two repetitions each of 15 s for flexion and extension). After the participants were fitted with the headgear and reflective markers, they were seated within the external force applicator (Fig. 2) and instructed to sit with their head up and eyes facing forward, looking straight ahead. The thorax and pelvis of the participants were positioned against the back of a chair and stabilized with a Velcro (Velcro U.S., Manchester, NH) strap to minimize extraneous body movements (24). Video was also used to ensure no trunk movement during the trials. Participants performed three flexion and extension maximum voluntary isometric contractions. Participants then had their eyes and ears covered with modified goggles and earphones (24), respectively, to ensure no visual or auditory feedback during testing. The external force applicator’s pulley was then attached to the headgear and the 1-kg mass was placed on the landing surface for 10 s to allow participants to become accustomed to the amount of the external load. The 1-kg mass was then dropped 15 cm to apply the external load to the head–neck segment (Fig. 1) causing forced flexion (i.e., trapezius I eccentric tension) or forced extension (i.e., SCM eccentric tension). The amount of force applied was not normalized to body weight to simulate real life situations (e.g., soccer heading) where forces applied to the head are not related to body weight.

Trials were performed with and without participants’ knowledge of force application to simulate sport situations. Three known force application trials occurred first followed by three unknown force application trials for each direction. Direction of force application was randomized. During trials with participants’ knowledge, they were instructed to prepare (preactivate) their neck muscles for the external force application (24,31). There was a 3-s count down and then the mass was dropped. During trials without participant knowledge, participants were instructed to relax their neck muscles and then resist as soon as they felt a tug (24,31). They were told that the force would be applied at some point during the next 30 s. The duration before the drops was random and ranged from 5 to 25 s. Also, unknown trials were removed if muscle preactivation occurred. The average of three trials was used as the criterion measure for all analyses.

Data Analysis

Data were analyzed using descriptive and inferential statistics. Statistical analyses consisted of multiple multivariate and univariate analyses of variance, follow-up univariate analyses of variance and t-tests with Bonferroni correction. Alpha level was set at P < 0.05.

Potential covariates in the kinematic, EMG, head–neck segment stiffness, and isometric strength analyses were head–neck segment mass (kg) and length (cm) as well as neck girth (cm). Covariates were correlated a priori with the appropriate dependent variables. A correlation value r ≥ 0.60 was used as the criterion for inclusion as a covariate. Neck girth and head mass were correlated (r > 0.60) with head–neck segment isometric flexion and extension strength. They were not statistically significant (P > 0.05) in the ANCOVA model, however, and were therefore not included in the analysis. None of the potential covariates met this criterion for any other dependent variable and were not utilized. The SPSS for Windows, Version 11.5, statistical program (SPSS, Inc., Chicago, IL) was used for data analysis.

RESULTS

Head–neck segment anthropometric assessments.

t-tests revealed significant gender differences in head–neck segment mass and neck girth (Table 1). Females exhibited 43 and 30% less head–neck segment mass and neck girth than males, respectively. There were no significant gender differences in head–neck segment length.

TABLE 1
TABLE 1:
Participant anthropometric measurements by gender.

Head–neck segment kinematics.

A 2 × 2 × 2 MANOVA revealed a significant gender × force application interaction effect (Table 2). The follow-up ANOVA revealed a significant interaction for angular acceleration (F(1,37) = 4.72, P = 0.036) but not for angular displacement (F(1,37) = 1.15, P = 0.290). Post hoc analysis for the two-factor interaction revealed a significant difference (t(1,38) = −2.8, P = 0.007) between the known and unknown force applications for males. Specifically, males exhibited 25% less angular acceleration during the known trials versus the unknown trials. There was no significant difference (t(1,36) = −0.28, P = 0.778) in females between the known and unknown force application conditions. Also, there were significant differences between genders during the known (t(1,37) = 5.97, P < 0.001) and unknown (t(1,37) = 3.3, P = 0.001) force applications. Specifically, females had 70 and 31% more head–neck segment angular acceleration than males during the known and unknown conditions, respectively.

TABLE 2
TABLE 2:
Means and standard deviations for head–neck segment kinematic data.

The 2 × 2 × 2 MANOVA also revealed a significant gender main effect, and the follow-up ANOVA indicated a main effect for angular displacement (F(1,37) = 43.02, P < 0.001). Results of the gender main effect revealed that head–neck segment angular displacement was 39% greater in females compared with males.

Peak muscle activity and muscle activity area.

A 2 × 2 × 2 MANOVA revealed no significant interaction effects but did show a significant main effect for gender (Table 3). The follow-up ANOVA revealed a significant gender effect for peak muscle activity (F(1,34) = 10.22, P = 0.003) and muscle activity area (F(1,34) = 14.28, P < 0.001) only. Females exhibited 79% more peak muscle activity and 117% more muscle activity area compared with males.

TABLE 3
TABLE 3:
Means and standard deviations for head–neck segment muscle activity.

Muscle onset latency.

A 2 × 2 ANOVA revealed a significant gender × force direction interaction (Table 4). Post hoc independent t-tests indicated a significant difference (t(1,36) = −3.32, P = 0.002) between genders during forced extension but not during forced flexion (t(1,36) = 0.692, P = 0.494). Muscle onset latency was 29% faster for females versus males for the Sternocleidomastiod and only 9% faster in the trapezius.

TABLE 4
TABLE 4:
Means and standard deviations for muscle onset latency.

Head–neck segment stiffness.

A 2 × 2 × 2 ANOVA revealed a significant main effect for gender (Table 5). The significant gender main effect indicated that the females exhibited 29% less stiffness than the males.

TABLE 5
TABLE 5:
Means and standard deviations for head–neck segment stiffness.

Isometric head–neck segment muscle strength.

A 2 × 2 ANOVA revealed significant main effects for gender (Table 6). Females exhibited 49% less isometric neck muscle strength than males. Participants exhibited 30% more isometric strength during forced flexion than during forced extension.

TABLE 6
TABLE 6:
Means and standard deviations for head–neck segment flexor and extensor isometric strength.

DISCUSSION

The results of this study revealed gender differences in head–neck segment dynamic stabilization during head acceleration in response to an external force application. Females exhibited significantly greater head–neck segment peak angular acceleration and displacement than males despite initiating muscle activity earlier (SCM only) and using a greater percentage of their maximum head–neck segment muscle activity. In addition, knowledge of force application had no effect in limiting head–neck segment angular acceleration for females, whereas it resulted in a significant reduction for males. These findings are similar to that of previous research involving low speed rear-end car collisions (15) and indicate gender differences in the ability of physically active individuals to use their dynamic stabilizers for protection against head injury. The reason for the greater head–neck segment angular acceleration in females may be related to their lower levels of strength, neck girth, and head mass, resulting in less head–neck segment stiffness compared with males.

Females exhibited 29% less stiffness and almost 50% less isometric strength than males, regardless of force direction. Previous researchers reported similar stiffness (24) and isometric strength (9,25) gender differences, which they attributed to the amount of muscle tissue. In the present study, females had 23% less neck girth than males. Because more tissue correlates positively with greater joint resistance to motion (6), females should perform head–neck segment resistance training to increase neck girth.

In the present study, female head–neck segment mass was 43% less than males and may have contributed to the significant gender differences in acceleration. Risk of concussion is directly related to the amount of head acceleration (linear and angular) during force application (10). Following Newton’s law of acceleration, for a given force application less head mass correlates with greater head acceleration (16) and risk of concussion (1,26). Schneider and Zernicke (26) reported an increased risk of head injury during soccer headings for individuals with a lower head mass to soccer ball mass ratio (i.e., women and children).

Force application knowledge trials were not randomized in this study (known then unknown trials) in order to enhance protection for the subjects during this novel task involving the head and neck. This design could have caused a learning effect allowing subjects to be more prepared for the unknown force application trials. We believe this effect was reduced because any unknown trials with muscle preactivity were not included in any analysis. Also, this design elicited a significant gender by force application knowledge interaction for angular acceleration as well as main effects (EMG and stiffness not discussed).

The gender differences in kinematics cannot be attributed to improper muscle function. The EMG data indicated that females had significantly greater muscle activity than males regardless of force application or force direction. Because EMG data were collected over a 250-ms window beyond initiation of force application, this indicates a combination of greater reflex and voluntary muscle activities (13). Greater head angular acceleration and displacement should stimulate a higher percentage of mechanoreceptors in the muscular (i.e., muscle spindle) and articular tissues (i.e., paciniform corpuscles) of the cervical spine leading to a greater activation of motor neurons and more reflex muscle firing (7,14).

Females also had a significantly faster onset latency time in the sternocleidomastiod than males during the unknown condition. This result is similar to previous research (4) and may be related to the greater acceleration values for females during forced extension. The unknown forced extension trials for females elicited the greatest peak angular accelerations. Greater angular acceleration indicates a faster initial movement of the head after force application. Head movement is monitored by the vestibular apparatus of the inner ear (13) and causes a vestibulocollic reflex to the appropriate stabilizing head–neck segment muscles (e.g., SCM). The present study results indicate that the gender and condition with the greatest amount of head–neck angular acceleration yielded the fastest muscle reflex.

Brault et al. (4) reported significantly faster SCM muscle onset latency time in females compared with males during low-speed rear-end collision testing but no gender differences in peak SCM muscle activity. The present study’s results indicated faster SCM onset latency times and a larger amount of muscle activity for females. The difference in results of the two studies may be related to method of force application. Our subjects had to stabilize their heads with an added load applied directly to it, whereas the previous studies subjects needed only to control their own head mass. Therefore, time of muscle firing would be similar, yet more muscle activity would be required of our subjects.

The muscle onset latency times (SCM reflex times of 40 ms) were longer compared with those of previous research (13). The differences are attributed to the method of assessing the reflex. Ito et al. (13) reported SCM reflex times of 24 ms but tested participants positioned supine with their heads supported in a sling and their neck muscles relaxed. Their heads were then released causing an immediate acceleration due to gravity. In the present study, force was applied via a pulley attached to headgear that was worn by participants seated in a chair. The time from force application to head movement, when loading of the pulley occurred, would have resulted in the longer muscle onset latency time.

The amount of force used (and resultant angular accelerations elicited) in this study was by design lower than what is thought to cause a concussion (10). However, with the relatively low applied force gender differences in head–neck segment angular acceleration were still evoked. These differences may be because females exhibited significantly less head–neck segment isometric strength, neck girth, and head mass, resulting in lower levels of stiffness. This is important because angular acceleration is directly related to concussion (10). Although these subjects were not athletes the results suggest that females may be at greater risk of concussion in sports with greater loads being applied to the head (e.g., soccer, > 180 N) (2). Future research should examine other populations and the effect of head–neck resistance training on kinematic and EMG responses to force application for the purpose of concussion prevention.

CONCLUSIONS

For our subject demographic, the results of this study revealed gender differences in head–neck segment dynamic stabilization during head acceleration in response to an external force application. Females exhibited significantly greater head–neck segment peak angular acceleration and displacement than males despite initiating muscle activity earlier (SCM only) and using a greater percentage of their maximum head–neck segment muscle activity.

ACKNOWLEDGMENTS

We would like to thank Jack Reed for his help in the construction of the external force applicator.

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Keywords:

HEAD INJURY; CONCUSSION; BIOMECHANICS; NEUROMUSCULAR CONTROL; PATHOMECHANICS

©2005The American College of Sports Medicine