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IRCOBI Conference – Graz (Austria) September 2004 269
EFFECTS OF REFLEXIVE CERVICAL MUSCLE CONTRACTION ON WHIPLASH
KINEMATICS AND INJURY
Brian D. Stemper, Narayan Yoganandan, Frank A. Pintar, Thomas A. Gennarelli
Department of Neurosurgery, Medical College of Wisconsin, and Zablocki VA Medical Center
Milwaukee, WI, USA
ABSTRACT
The ability of neck muscles to react in time to mitigate whiplash injury in unaware occupants
remains unclear. A validated head-neck computer model was implemented to determine the effect of
reflexive contraction on capsular ligament elongations in whiplash. Contraction parameters were
chosen for the maximum kinematic effect. Contraction decreased segmental angulations less than
10% during the retraction phase and decreased capsular ligament elongations in the middle while
slightly increasing elongations at the inferior and superior ends of the cervical spine. Due to inherent
reflexive contraction delays, it remains unlikely that reflexive contraction can significantly affect the
whiplash injury mechanism.
Key Words: Biomechanics, rear impacts, spine, computer model, facet joints
THE ABILITY OF NECK MUSCLES to mitigate whiplash injury in the unaware occupant continues
to be a topic of debate in biomechanics literature. By measuring electromyographical (EMG) signals
of superficial neck muscles, experimental investigators concluded that reflexive contraction of the
neck muscles (in particular the sternocleidomastoid) can affect spinal kinematics in whiplash
(Magnusson et al., 1999; Brault et al., 2000; Kumar et al., 2002). These assumptions are largely based
on the relative timing between EMG and head or thorax accelerations. However, a separate group of
researchers assert that whiplash injury in the unaware occupant occurs during the initial stages of
head-neck kinematics (retraction) prior to the generation of significant muscle forces (Geigl et al.,
1994; Ono et al., 1997; Deng et al., 2000; Stemper et al., 2003). The most important issue in the
present debate is the timing of the whiplash injury mechanism relative to the timing of reflexive
muscle contraction and force generation. To affect the whiplash injury mechanism, cervical muscles
must react and generate sufficient forces to decrease spinal kinematics prior to the time of injury
occurrence.
The present investigation exercised a validated head-neck computer model in whiplash to
measure region-dependent elongation of facet joint capsular ligaments. To test the effects of muscle
contraction in the unaware occupant, the model implemented reflexive muscle contraction with
parameters obtained from literature. Contraction parameters were chosen to provide the greatest
kinematic effect. The hypothesis was that reflexive muscle contraction would not markedly alter
ligament elongations during the retraction phase because of inherent physiological timing delays
associated with this type of contraction.
LITERATURE REVIEW
Experimental investigators implemented human volunteers, full-body cadavers, and intact
head-neck complexes to correlate neck motions to potential whiplash injury mechanisms (Geigl et al.,
1994; Ono et al., 1997; Deng et al., 2000; Cusick et al., 2001; Kumar et al., 2002). Three phases of
cervical spinal kinematics result from whiplash loading (Ono et al., 1997; Davidsson et al., 1998; Ono
et al., 1999). During the initial phase, the thorax is displaced anteriorly due to interaction with the
seatback, while the head remains stationary due to its inertia, resulting in retraction of the head relative
IRCOBI Conference – Graz (Austria) September 2004 270
to the thorax. To compensate for head retraction, the cervical spine sustains non-physiologic S-
curvature, characterized by extension in lower and flexion in upper segments (Cusick et al., 2001). In
the second phase, loading is transferred up the cervical spine, and the entire head-neck complex moves
into extension and continues in that mode until the head contacts the head restraint. In the final phase,
the head rebounds from the head restraint, forcing the head-neck complex into flexion. Injury
mechanisms hypothesized to occur in each of these stages are discussed below.
Reflexive contraction of the neck muscles can be modeled in a unique sequence that correlates
to experimental findings. The three specific stages of reflexive muscle contraction are reflex delay,
electromechanical delay (EMD), and build-up of muscle forces (Siegmund and Brault 2000). Reflex
delay is most commonly defined in human volunteer rear impact studies as the delay between the
stimulus (impact) and initiation of electrical muscle activity (Szabo and Welcher 1996; Ono et al.,
1997; Kaneoka et al., 1999; Magnusson et al., 1999; Brault et al., 2000; Kumar et al., 2002; Siegmund
et al., 2003). These investigations reported mean EMG reflex delays from the initiation of T1
acceleration to be 48 msec. The second phase, EMD, has received considerable attention in literature,
and is characterized by the conversion of electrical muscle activity into mechanical force (Table 1).
This time delay varies depending on type of contraction, velocity of movement, involved muscle,
fatigue, age, and gender (Norman and Komi 1979; Zhou et al., 1995). EMD values for reflexive
contractions are typically shorter than voluntary contractions (Table 1). The final phase involves the
build-up of tensile force to the maximum value for the neural activation level, characterized by muscle
force rise time. This value has been reported to be 81 msec for cervical muscles in canines subjected
to 3 to 5 g vertical accelerations (Tennyson et al., 1977). Although literature demonstrated a range of
values for the stages of muscle contraction, the earliest time of maximum muscle contraction is
approximately 140 msec. The ability of reflexive neck muscle contraction to mitigate whiplash injury
largely depends on the kinematic phase during which injury occurs. Muscle contraction likely has a
significant affect on kinematics occurring later than 140 msec. The effect of this contraction prior to
140 msec is however, unclear. This is an objective of the present study.
Table 1: Times of electromechanical delay reported in the literature.
Author Muscle Contraction Type EMD (ms)
(Cavanagh and Komi 1979) biceps and brachioradialis voluntary concentric 56
(Corcos et al., 1992) biceps voluntary isometric 13-31
(Granata et al., 2000) knee extensor reflex (isometric) 54 ± 15
(Granata et al., 2004) lumbar paraspinal reflex 29-34
(Nilsson et al., 1977) vastus lateralis voluntary concentric 95 ± 3
(Norman and Komi 1979) biceps and triceps voluntary eccentric 26-41
(Vint et al., 2001) biceps voluntary concentric 84 ± 13
(Vos et al., 1991) vastus lateralis voluntary concentric 82 ± 15
(Zhou et al., 1995) knee extensor reflex (isometric) 22-24
Soft tissue injuries to the cervical spine resulting from automotive rear impacts were first
recognized in 1928 (Crowe 1928). By 1995, 10,000 clinical and experimental articles were published
on the subject (Spitzer et al., 1995). During that time, research with clinical patients, human
volunteers, cadavers, animals, and crash test dummies led to a number of whiplash injury theories.
One of the first theories was injury to the anterior cervical structures as a result of hyperextension of
the head-neck complex (Macnab 1971). This theory was tested using human volunteers, monkeys,
and crash test dummies. To prevent hyperextension injuries in automotive rear impacts, in 1969 the
United States government mandated head restraints for all new passenger cars within its borders
(NHTSA 1969). However, these safety devices were only marginally successful at mitigating
whiplash injury, decreasing overall injury risk in rear impacts by 5 to 20% (O'Neill et al., 1972;
Kahane 1982). Other injury theories included neuronal degeneration due to pressure gradients in the
spinal canal resulting from differential motion between the head and thorax during the retraction phase
(Svensson et al., 1993; Schmitt et al., 2003), and injury to the anterior cervical muscles due to
IRCOBI Conference – Graz (Austria) September 2004 271
eccentric contraction during the retraction phase (Brault et al., 2000). Both theories recognized the
significance of initial head retraction resulting in abnormal S-curvature in the cervical spine.
Recently, a number of clinical and experimental investigations focused on injury to cervical facet
joints as a likely mechanism of injury (Barnsley et al., 1995; Ono et al., 1997; Yang et al., 1997;
Kaneoka et al., 1999; Cusick et al., 2001; Yoganandan et al., 2001). Lower cervical facet joints
sustained tension in anterior regions and compression in posterior regions during the retraction phase
(Cusick et al., 2001). Excess tension in the joint capsular ligament can lead to subcatastrophic or
catastrophic tissue failure (Yoganandan et al., 1989; Winkelstein et al., 2000; Siegmund et al., 2001).
Clinical studies correlated injury to lower cervical facet joints with the most commonly reported
whiplash symptoms (Barnsley et al., 1995). Because of these clinical and experimental investigations,
the injury mechanism examined in the present study was tensile injury to the facet joint capsular
ligaments during the retraction phase, the time of nonphysiologic cervical S-curvature, and as
described in the previous paragraph this study was focused on analyzing the effect of reflexive
contraction in the unaware occupant.
METHODS
A head-neck computer model was used to investigate the effects of reflexive muscle
contraction on kinematics of the cervical spine in whiplash. The model was exercised using
MADYMO software (TNO, the Netherlands) and consisted of the head, seven cervical vertebrae, first
thoracic vertebra, and soft tissues of the cervical spine incorporated as ligaments, intervertebral discs,
facet joints, and passive neck musculature. Bony components of the head-neck complex were
modeled as rigid bodies, including mass and inertial properties of bone and surrounding soft tissues
(de Jager 1996; van der Horst 2002). Bony geometry was incorporated using CT images of a single
male specimen (de Jager 1996). Level-dependent, nonlinear, viscoelastic soft tissue material
properties were based on literature (Table 2). Facet joints were modeled with high compressive
stiffness and zero shear and tensile stiffness, according to synovial joint mechanics. Resistance to
facet joint tension and shear motion was accomplished through elongation of the capsular ligaments.
Table 2: Material properties of the head-neck computer model.
Spinal component Spinal level Loading Material Property Reference
Upper cervical Anterior longitudinal ligament
Lower cervical Tension (Pintar 1986; Yoganandan et al., 1989)
Upper cervical Posterior longitudinal ligament
Lower cervical Tension (Pintar 1986; Yoganandan et al., 1989)
Upper cervical Ligamentum flavum
Lower cervical Tension (Pintar 1986; Yoganandan et al., 1989)
Upper cervical Interspinous ligament
Lower cervical Tension (Pintar 1986; Yoganandan et al., 1989)
Upper cervical Capsular ligament
Lower cervical Tension (Pintar 1986; Yoganandan et al., 1989)
Shear (Moroney et al., 1988)
Tension (Pintar et al., 1986) Intervertebral disc All levels
Compression (Eberlein et al., 1999)
Neck musculature was modeled using the MADYMO Hill-type muscle model (TNO
Automotive). The Hill model consists of parallel elastic and contractile elements in series with two
elastic elements. The contractile element controls the active force generation of the muscle, the elastic
elements account for stiffness of the muscle fibers, surrounding tissue, and the tendons and
aponeurosis. Muscles were attached to individual vertebrae according to local coordinates, creating
‘sliding’ interfaces that permitted localized muscle elongation and allowed muscles to ‘wrap’ around
the vertebral column and develop more realistic lines of muscle action. In most cases, attachment
points were defined at each cervical level between origin and insertion. Passive muscle resistance to
IRCOBI Conference – Graz (Austria) September 2004 272
motion of the head-neck complex was approximated according to literature (Deng and Goldsmith
1987).
Muscles were divided into three activation groups by their kinematic effect: flexors, extensors,
and sternocleidomastoid. Because the sternocleidomastoid cannot be classified uniquely as a flexor or
extensor due to proximal attachment points anterior to the cervical spine and distal attachment points
posterior to the occipital condyles, this muscle was included in its own kinematic activation group.
Reflexive muscle contraction was divided into three phases consisting of the reflex delay, EMD, and
muscle rise time (Figure 1). Muscles sustained zero neural activation (force generation) during the
first two phases and linearly ramped up to maximum activation during the muscle rise time. Two
reflex delays, 45 and 54 msec, were implemented to represent the literature (Ono et al., 1997;
Magnusson et al., 1999). Electromechanical delay was 13 msec (Corcos et al., 1992). Muscle rise
time to maximum neural activation lasted 81 msec (Tennyson et al., 1977). Muscle activation was
accomplished according to synergistic assumptions, wherein all muscles of a specific group were
activated to the same level. The ratio of maximum activation levels between the three muscle groups
was determined prior to whiplash simulations in order to obtain a sagitally balanced contraction to
balance the flexion and extension moments applied to the head. It was determined that 90% activation
of the flexors, 45% of the sternocleidomastoid, and 30% of the extensors resulted in +1g vertical
acceleration of the head to counterbalance the acceleration of gravity, with minimal sagittal plane
rotation of the head. These activation levels were implemented during whiplash simulations.
0
100
0 200time (msec)
{
EMD
Reflex
Time
Mu
scle
Forc
e R
ise T
ime
Maxim um Neura l
Act ivat ion
(45 to 54 msec)
(13 msec)
(81
mse
c)
0
100
0 200time (msec)
{
EMD
Reflex
Time
Mu
scle
Forc
e R
ise T
ime
Maxim um Neura l
Act ivat ion
(45 to 54 msec)
(13 msec)
(81
mse
c)
Figure 1: Neural activation (percentage of maximum activation) of the flexor muscle group.
Whiplash loading was imparted to the model by accelerating the T1 vertebra anteriorly. The
acceleration pulse was integrated to compute T1 change in velocity. Rear impact velocity of 2.6 m/sec
was imparted for all simulations. Prior to initiation of T1 acceleration, the occipital condyles were
positioned directly superior to the T1 vertebral body, and the Frankfort plane was maintained
horizontal. The cervical spine demonstrated normal lordotic posture, and the T1 vertebra was given an
anterior orientation of 25 deg to simulate normal driving alignment. During the entire whiplash
simulation, the T1 vertebra was constrained against rotation and superior and lateral translation.
Overall, level-by-level segmental, and facet joint kinematic responses without muscle contraction had
been previously validated with respect to experimental head-neck cadaver specimens subjected to
similar rear impact loading magnitudes (Stemper et al., 2004). The validation process is discussed in
further detail below.
To quantify the effects of reflexive muscle contraction on kinematics of the head-neck
complex, facet joint capsular ligament elongations at the C2-C3 through C6-C7 levels were compared
between the two muscle contraction simulations (54 and 45 msec reflex delays) and the simulation
without muscle contraction. Ligament elongation was defined as the increase in length from the
original ligament length at the initiation of T1 acceleration. In particular, facet joint capsular ligament
elongations were computed in four anatomic regions of the joint (ventral, lateral, dorsal, and medial)
during the time of maximum cervical S-curvature. This time was determined as the time of maximum
segmental flexion at the C2-C3 level during the retraction phase.
IRCOBI Conference – Graz (Austria) September 2004 273
RESULTS
The model demonstrated retraction and extension phases of whiplash kinematics for the
simulations with and without reflexive muscle contraction (Figure 2). Rebound did not occur due to
the absence of a head rest. Maximum S-curvature during the retraction phase occurred earlier with
decreasing reflex delays: 76 msec for the simulation without contraction, and 75 and 71 msec for the
54- and 45-msec reflex delay simulations, respectively. Reflexive muscle contraction also resulted in
a more timely transition from retraction to extension phases, occurring at 99 msec for the 54-msec
reflex delay and 94 msec for the 45-msec reflex delay, compared to 114 msec for the simulation
without contraction. Reflexive muscle contraction decreased the overall head extension angle relative
to T1 at the time of maximum S-curvature by 5.1% for the 54-msec delay and 24.9% for the 45-msec
delay.
Retraction Extension Figure 2: Initial position, retraction, and extension phases.
Neural activation of neck muscles initiated at 67 and 58 msec for the 54- and 45-msec reflex
delays. Muscle groups attained sub-maximal activation levels during the time of maximum S-
curvature (Table 3). Maximum activation, which coincided with maximum tensile force generation,
occurred at 148 and 139 msec.
Table 3: Percent of maximum neural activation during maximum S-curvature.
Flexors Extensors SCM
45-msec delay 16.0 16.0 16.0
54-msec delay 9.9 9.9 9.9
No contraction 0.0 0.0 0.0
During maximum S-curvature, flexion at the C2-C3 segmental level and extension at the C3-
C4 through C6-C7 levels was evident (Figure 3). Reflexive muscle contraction decreased segmental
angulations at all cervical levels, with the 45-msec reflex delay having a greater kinematic effect. The
54-msec reflex delay had a minimal effect on segmental angulations at the time of maximum S-
curvature (<1%). At the C2-C3, C4-C5, C5-C6, and C6-C7 levels, the 45-msec reflex delay
contraction decreased segmental angulations by less than 10%.
IRCOBI Conference – Graz (Austria) September 2004 274
-5
0
5
10
C2-C3 C3-C4 C4-C5 C5-C6 C6-C7
Seg
men
tal
angle
(deg
)
No con t ract ion
54-msec delay
45-msec delay
Figure 3: Segmental angulations at the time of maximum S-curvature.
Facet joint capsular ligaments demonstrated region- and level-dependent elongations during
the whiplash simulation without muscle contraction (Figure 4). Elongation magnitudes were below
2.0 mm in all anatomic regions. At the C2-C3 level, the dorsal anatomic region of the capsular
ligament sustained maximum elongation. At the C3-C4 through C6-C7 levels, lateral anatomic
regions of the capsular ligament sustained maximum elongations. Ligament elongations in the lateral
joint regions increased inferiorly, with maximum elongation at the C6-C7 level.
0
0.5
1
1.5
2
2.5
C2-C3 C3-C4 C4-C5 C5-C6 C6-C7
Lig
am
ent
elo
ng
ati
on
(m
m)
Ventral
Lateral
Dorsal
Medial
D L L L L
Figure 4: Capsular ligament elongations for the simulation without muscle contraction.
Simulations with reflexive muscle contraction demonstrated maximum elongations in the
same joint regions as the simulation without muscle contraction (C2-C3: dorsal, C3-C4 to C6-C7:
lateral). Facet joint capsular ligament elongations in those regions were compared to the simulation
without contraction (Figure 5). Muscle contraction decreased capsular ligament elongations at the C3-
C4 and C4-C5 levels by less than 16 percent for the 54-msec delay, and 16.4 percent at the C4-C5
level and 32.7 percent at the C3-C4 levels for the 45-msec delay. However, reflexive contraction
increased elongations at the C2-C3, C5-C6, and C6-C7 levels by a maximum of 21 percent.
IRCOBI Conference – Graz (Austria) September 2004 275
0.0
0.5
1.0
1.5
2.0
2.5
3.0
C2-C3 C3-C4 C4-C5 C5-C6 C6-C7
Lig
amen
t el
ongat
ion (
mm
)
No con t ract ion
54-msec delay
45-msec delay
DORSAL LATERAL LATERAL LATERAL LATERAL
Figure 5: Capsular ligament elongations at the time of maximum S-curvature.
DISCUSSION
Stabilization of the head-neck complex in an unaware occupant is accomplished in three
stages: passive structures (spinal soft tissues, passive musculature) act first, followed by reflexive
muscle contraction, and finally, voluntary muscle contraction (Simoneau et al., 2003). In a whiplash
event, wherein the injury mechanism likely occurs prior to head restraint contact, kinematics are
dominated by the first two stages. The ability of the occupant to mitigate injurious loading largely
depends on the mechanical properties of the passive structures and the timing of the reflexive muscle
contraction. Low-velocity whiplash loading has been experimentally shown to exceed the mechanical
thresholds of the passive neck structures, resulting in soft tissue spinal injury in human cadavers
subjected to whiplash (Deng et al., 2000; Yoganandan et al., 2000; Yoganandan et al., 2001).
However, the debate over the ability of reflexive neck muscle contraction in the unaware occupant to
mitigate whiplash injury is unresolved. Human cadaver studies cannot resolve this issue, and
awareness effects limit the applicability of human volunteer studies. The purpose of the present
investigation was to quantify the timing of reflexive neck muscle contraction relative to kinematics of
the facet joint capsular ligaments, structures clinically and experimentally linked to whiplash injury.
The present investigation focused on the time of injury as the non-physiologic S-curvature that occurs
during the retraction phase of whiplash kinematics, prior to head restraint contact.
HEAD RESTRAINTS
The effectiveness of head restraints in mitigating whiplash injury was reported to be minimal
(O'Neill et al., 1972; Kahane 1982). This finding implies that whiplash injury likely occurs prior to
the time that the head contacts the head restraint, for conventional restraints positioned with a finite
backset. A survey of the experimental rear impact literature revealed that the average time of head
restraint contact after initiation of T1 acceleration is 77.9 msec (Table 4). In the present study,
maximum S-curvature of the cervical spine occurred just prior to the mean time of head restraint
contact. During the cervical S-curvature, reflexive muscle contraction activation levels were less than
20% of maximum levels (Table 3). The kinematic effect of reflexive muscle contraction was also
minimal during this time, decreasing the overall head to T1 angulation by less than 25% and level-by-
level segmental angulations by less than 10% at the shortest reflex delay. Present results lead to the
conclusion that reflexive muscle contraction is unlikely to mitigate whiplash injuries occurring during
the retraction phase due to inherent delays involved in the initiation and buildup of these contractions.
IRCOBI Conference – Graz (Austria) September 2004 276
Table 4: Experimental times of head restraint contact.
Investigator Impact velocity
(m/sec)
T1 acceleration
(msec)
HR contact
(msec)
T1 acc to HR contact
(msec)
(Davidsson et al., 1998) 1.9 50 98 48
(Deng et al., 2000) 2.0 120 230 110
(Deng et al., 2000) 2.2 140 266 126
(Hell and Langwieder 1998) 2.6 50 130 80
(Hell and Langwieder 1998) 2.6 40 110 70
(McConnell et al., 1993) 2.2 100 140 40
(Siegmund et al., 1997) 1.1 30 118 88
(Siegmund et al., 1997) 2.2 30 94 64
(Kroonenberg et al., 1998) 1.8 - 2.6 0 75 75
Mean 77.9
CAPSULAR LIGAMENT ELONGATIONS
Reflexive muscle contraction had a marginal effect on capsular ligament elongations.
Maximum ligament elongations occurred in the same anatomic facet joint regions during both
simulations with reflexive muscle contraction and the simulation without muscle contraction. Muscle
contraction had an inconsistent effect on the magnitude of capsular ligament elongations. In the
middle cervical spine, contraction decreased elongations by only 16.4 percent at the C4-C5 level.
However, reflexive contraction increased elongations at the inferior and superior ends of the cervical
spine up to 21 percent. Because experimental whiplash studies produced injuries in the cervical facet
joints (Deng et al., 2000; Yoganandan et al., 2000) and due to clinical findings linking facet joint
injury with common whiplash symptoms (Barnsley et al., 1995), cervical facet joints are strongly
implicated in whiplash injury. Although the effect of reflexive contraction on capsular ligament
elongations was relatively small, present results suggest that reflexive muscle contraction may slightly
decrease the likelihood of whiplash injury at specific cervical levels, while slightly increasing the
likelihood of injury at other levels. The dependence of these results on reflex delay, with a greater
kinematic effect for shorter delays, indicates that shorter reflex delays than those used in the present
study would have a larger effect on spinal kinematics. However, shorter contraction delays are
unlikely in vivo as the present study maximized the kinematic effect of reflexive contraction according
to literature.
REFLEXIVE CONTRACTION PARAMETERS
Reflexive muscle contraction parameters were chosen to provide the maximum effect on
whiplash kinematics. The 45-msec reflex delay, although not the minimum value reported in
literature, was the shortest reflex delay obtained for human volunteers subjected to similar magnitudes
of rear impact loading (Ono et al., 1997). Because of a reported correlation between impact severity
and reflex delay (Kumar et al., 2002), the 54- and 45-msec delays were chosen to represent literature.
Likewise, the 13-msec EMD was chosen to minimize this delay and maximize the effect of reflexive
contraction on spinal kinematics. EMD values were reported as high as 95 msec for voluntary
contractions (Nilsson et al., 1977), and 54 msec for reflexive contractions (Granata et al., 2000).
However, the 13-msec delay, although obtained in voluntary contraction, was reported to be a more
accurate measurement (Siegmund and Brault 2000) and was, therefore, chosen to maximize the
kinematic effect. Muscle rise time to maximum neural activation (81 msec) was also consistent with
literature (Tennyson et al., 1977; Szabo and Welcher 1996; Magnusson et al., 1999; Kumar et al.,
2002). In the present study, maximum neural activation corresponded to maximum muscle force
generation. For all neck muscles, mean muscle force rise times obtained in human volunteer and
animal studies were reported between 63 msec (Szabo and Welcher 1996) and 253 msec (Kumar et al.,
2002). An investigation into the sensitivity of spinal kinematics to muscle rise time conducted prior to
this study revealed that segmental angulations changed by less than 1.3 percent at the C2-C3 and C4-
C5 through C6-C7 levels and by 10.4 percent at the C3-C4 level with variation of the muscle rise time
IRCOBI Conference – Graz (Austria) September 2004 277
between 63 and 81 msec. The most sensitive contraction parameter, as indicated by the results of the
present study, was reflex delay. Minimizing the time to initiation of neural activation through reflex
delay and electromechanical delay resulted in the maximum effect of reflexive muscle contraction on
spinal kinematics.
MODEL VALIDATION
The present computer model was validated with respect to the overall head to T1 angulations,
level-by-level segmental angulations (C2-C3 through C6-C7 levels), and region-dependent facet joint
motions (C4-C5 through C6-C7 levels) obtained from ten intact head-neck cadaver specimens
subjected to 1.8 and 2.6 m/sec rear impacts (Stemper et al., 2004). Experimental specimens were
subject to identical boundary conditions as the computer model (e.g., initial head and spinal
orientations and T1 constraints). Validation corridors were developed from experimental specimens
consisting of the mean plus and minus one standard deviation kinematic responses. Corridors were
mass-scaled to account for biological variation in specimen anthropometry according to accepted
procedures (Maltese et al., 2002). Computer model response falling within the corridors was
considered valid. Validation plots for overall motion, segmental angulations from C2-C3 through C6-
C7 levels, and facet joint motions in the anterior and posterior joint regions from C4-C5 through C6-
C7 levels are presented in appendix A. While previous models used in the study of whiplash have
focused validation efforts primarily on head and thoracic accelerations, the present model is the first to
incorporate overall motions, level-by-level segmental angulations, and region-dependent facet joint
motions in the validation process.
INITIAL OCCUPANT POSITIONING
Initial positioning of the model in the present study focused on the “normal” automotive
occupant characteristics; lordotic posture, horizontal Frankfort plane, physiologic occipital condyle
positioning. This assumption was made for the sake of consistency and to simplify the analysis.
However, soft tissue injuries resulting from rear impacts can also occur under separate mechanisms
and with different factors. For example, the absence of an automotive head restraint may lead to
hyperextension injuries in the anterior cervical spine (i.e., anterior longitudinal ligament and endplate
failures). Out of position occupants may also exhibit separate pathologies. Axial rotation of the head
prior to impact places an additional pre-strain on the contralateral spinal components that may lower
the threshold of injury or alter the injury mechanism. For the sake of consistency and to model the
most typical whiplash event, occupant position was strictly controlled in the “normal” position prior to
impact in all simulations.
LIMITATIONS
A limitation of the present study was that of synergistic muscle contractions, wherein all
muscles of each kinematic group were activated to identical levels. Reflexive contraction of the neck
muscles in vivo likely occurs in a more complex sequence involving unique contraction levels, rates,
and timing for individual muscles (Winters and Stark 1988). While some of these parameters were
defined for the superficial muscles using EMG analysis of volunteers subjected to rear impacts (Szabo
and Welcher 1996; Ono et al., 1997; Kaneoka et al., 1999; Magnusson et al., 1999; Brault et al., 2000;
Kumar et al., 2002; Siegmund et al., 2003), data is clearly lacking for deep muscles. Because of this
inconsistency, it was necessary to assume equal contraction levels, rates, and timing for muscles with
similar kinematic effects.
A second limitation was the T1 constraint preventing rotation and superior and lateral
translation. This constraint was included because the model was validated with respect to an intact
head-neck cadaver model subject to the same constraints (Stemper et al., 2004) and because the
literature on thoracic ramping is not consistent. While it is well acknowledged that the upper thorax
sustains ramping motion in whiplash due to the interaction with the seatback, the magnitude and
timing of this motion as reported in full-body cadaver and human volunteer experiments is
IRCOBI Conference – Graz (Austria) September 2004 278
inconsistent (McConnell et al., 1993; Davidsson et al., 1998; Deng et al., 2000; Yoganandan et al.,
2000). This ramping motion is likely dependent upon a number of factors, including stiffness of the
seatback, curvature of the thoracic spine, angle of the seatback, magnitude of rear impact, occupant
awareness, and occupant posture. Experimental results quantify this variability, wherein mean T1
rotation at 100 ms for a seatback angle of 20 deg was twice the magnitude of that for a seatback angle
of 0 deg (Deng et al., 2000). To eliminate this variability, T1 was constrained in all degrees of
freedom except anterior displacement in the present analysis.
CONCLUSIONS
This study quantified the effects of reflexive muscle contraction on cervical spine kinematics
during the retraction phase. Kinematics of the cervical spine during this phase are such that abnormal
loading patterns are placed on the cervical facet joints that may result in catastrophic or
subcatastrophic injury to the capsular ligaments. A validated head-neck computer model was exposed
to 2.6 m/sec rear impacts and reflexive contraction was modeled using parameters obtained from
literature to produce the maximum kinematic effect. Results demonstrated that muscle contraction in
the unaware occupant has a minimal affect on segmental angulations during the retraction phase.
Kinematic response demonstrated a dependence upon reflex delay, with the shorter delay resulting in a
larger effect on capsular ligament elongations. However, the change in ligament elongation between
the shortest reflex delay and the simulation without contraction was less than 21 percent at the C2-C3,
C4-C5, C5-C6, and C6-C7 levels. Results for the present study demonstrated the importance of the
retraction phase of whiplash kinematics and indicated that reflexive muscle contraction in the unaware
occupant may play a secondary role in minimizing the likelihood of capsular ligament injury during
whiplash.
ACKNOWLEDGMENTS
This study was supported in part by PHS CDC Grant R49CCR-515433 and the Department of
Veterans Affairs Medical Research.
REFERENCES
Barnsley, L., Lord, S. M., Wallis, B. J. and Bogduk, N. “The prevalence of chronic cervical zygapophyseal joint
pain after whiplash.” Spine 20(1): (1995) 20-26.
Brault, J., Siegmund, G. and Wheeler, J. “Cervical muscle response during whiplash: Evidence of a lengthening
muscle contraction.” Clin Biomech 15: (2000) 426-435.
Cavanagh, P. R. and Komi, P. V. “Electromechanical delay in human skeletal muscle under concentric and
eccentric contractions.” Eur J App Phys 42: (1979) 159-163.
Corcos, D. M., Gottlieb, G. L., Latash, M. L., Almeida, G. L. and Agarwal, G. C. “Electromechanical delay: an
experimental artifact.” J Electromyogr Kinesiol 2(2): (1992) 59-68.
Crowe, H. E. Injuries of the cervical spine. Western Orthopaedic Association, (1928) San Francisco.
Cusick, J. F., Pintar, F. A. and Yoganandan, N. “Whiplash syndrome: Kinematic factors influencing pain
patterns.” Spine 26(11): (2001) 1252-1258.
Davidsson, J., Deutscher, C., Hell, W., Linder, A., Lovsund, P. and Svensson, M. Y. Human volunteer
kinematics in rear-end sled collisions. International Research Council on the Biomechanics of Impact
(IRCOBI), (1998) Goteborg, Sweden, 289-301.
de Jager, M. Mathematical head-neck models for acceleration impacts. The Netherlands, University of
Eindhoven: (1996).
Deng, B., Begeman, P. C., Yang, K. Y., King, A. I. and Tashman, S. “Kinematics of human cadaver cervical
spine during low speed rear-end impacts.” Stapp Car Crash Journal 44: (2000) 171-188.
Deng, Y. C. and Goldsmith, W. “Response of a human head/neck/upper-torso replica to dynamic loading--II.
Analytical/numerical model.” J Biomech 20(5): (1987) 487-497.
Eberlein, R., Frohlich, M. and Hasler, E. M. Finite-element analysis of intervertebral discs. European Conference
on Computation Mechanics, (1999).
Geigl, B. C., Steffen, H., Leinzinger, P., Muhlbauer, M. and Bauer, G. The movement of head and cervical spine
during rear-end impact. International Research Council on the Biomechanics of Impact (IRCOBI),
(1994) Lyon, France, 127-137.
IRCOBI Conference – Graz (Austria) September 2004 279
Granata, K., Ikeda, A. and Abel, M. “Electromechanical delay and reflex response in spastic cerebral palsy.”
Arch Phys Med Rehab 81: (2000) 888-894.
Granata, K., Slota, G. and Bennett, B. “Paraspinal reflex dynamics.” J Biomech 37: (2004) 241-247.
Hell, W. and Langwieder, K. Reported soft tissue neck injuries after rear-end car collisions. International
Research Council on the Biomechanics of Impact (IRCOBI), (1998) Goteborg, Sweden, 261-274.
Kahane, C. An evaluation of head restraints-federal motor vehicle safety standard 202. Springfield, VA, National
Technical Information Service, NHTSA: (1982).
Kaneoka, K., Ono, K., Inami, S. and Hayashi, K. “Motion analysis of cervical vertebrae during whiplash
loading.” Spine 24(8): (1999) 763-770.
Kumar, S., Narayan, Y. and Amell, T. “An electromyographic study of low-velocity rear-end impacts.” Spine
27(10): (2002) 1044-1055.
Macnab, I. “The "Whiplash Syndrome".” Orthop Clin North Am 2(2): (1971) 389-403.
Magnusson, M. L. et al.,. “Cervical electromyographic activity during low-speed rear impact.” Eur Spine J 8:
(1999) 118-125.
Maltese, M. R., Eppinger, R. H., Rhule, H. H., Donnelly, B. R., Pintar, F. A. and Yoganandan, N. “Response
corridors of human surrogates in lateral impacts.” Stapp Car Crash Journal 46: (2002) 321-351.
McConnell, W. E., Howard, R. P., Guzman, H. M., Bomar, J. B., Raddin, J. H., Benedict, J. V., Smith, H. L. and
Hatsell, C. P. Analysis of human test subject kinematic responses to low velocity rear end impacts. SAE
World Congress and Exposition, (1993) Detroit, MI, 21-30.
Moroney, S., Schultz, A., Miller, J. and Andersson, G. “Load-displacement properties of lower cervical spine
motion segments.” J Biomech 21(9): (1988) 769-779.
NHTSA. FMVSS No. 202. Washington, National Highway Transportation Safety Administration: (1969).
Nilsson, J., Tesch, P. and Thorstensson, A. “Fatigue and EMG of repeated fast voluntary contractions in man.”
Acta Physiol Scand 101: (1977) 194-198.
Norman, R. and Komi, P. “Electromechanical delay in skeletal muscle under normal movement conditions.”
Acta Physiol Scand 106: (1979) 241-248.
O'Neill, B., Haddon, W., Kelley, A. and Sorenson, W. “Automobile head restraints: Frequency of neck injuries
insurance claims in relation to the presence of head restraints.” Am J Public Health 62(3): (1972) 569-
573.
Ono, K., Inami, S., Kaneoka, K., Gotou, T., Kisanuki, Y., Sakuma, S. and Miki, K. Relationship between
localized spine deformation and cervical vertebral motions for low speed rear impacts using human
volunteers. International Research Council on the Biomechanics of Impact (IRCOBI), (1999) Sitges,
Spain, 149-164.
Ono, K., Kaneoka, K., Wittek, A. and Kajzer, J. Cervical injury mechanism based on the analysis of human
cervical vertebral motion and head-neck-torso kinematics during low speed rear impacts. 41st Stapp Car
Crash Conference, (1997) Lake Buena Vista, FL, 339-356.
Pintar, F. A. The biomechanics of spinal elements. Milwaukee, WI, Marquette University: (1986).
Pintar, F. A., Myklebust, J., Sances Jr, A. and Yoganandan, N. Biomechanical properties of the human
intervertebral disk in tension. ASME Adv Bioeng, (1986) New York, NY, 38-39.
Schmitt, K. U., Muser, M., Niederer, P. and Walz, F. “Pressure aberrations inside the spinal canal during rear-
end impact.” Pain Res Manag 8(2): (2003) 86-92.
Siegmund, G. P. and Brault, J. R. Role of cervical muscles during whiplash. Frontiers in Whiplash Trauma:
Clinical & Biomechanical. N. Yoganandan and F. Pintar. The Netherlands, IOS Press: (2000) 295-320.
Siegmund, G. P., King, D. J., Lawrence, J. M., Wheeler, J. B., Brault, J. R. and Smith, T. A. Head/neck
kinematic response of human subjects in low-speed rear-end collisions. 41st Stapp Car Crash
Conference, (1997) Lake Buena Vista, FL, 357-385.
Siegmund, G. P., Myers, B. S., Davis, M. B., Bhonet, H. T. and Winkelstein, B. A. “Mechanical evidence of
cervical facet capsule injury during whiplash: A cadaveric study using combined shear, compress, and
extension loading.” Spine 26(19): (2001) 2095-2101.
Siegmund, G. P., Sanderson, D. J., Myers, B. S. and Inglis, J. T. “Awareness affects the response of human
subjects exposed to a single whiplash-like pertubation.” Spine 28(7): (2003) 671-679.
Simoneau, M., Tinker, S. W., Hain, T. C. and Lee, W. A. “Effects of predictive mechanisms on head stability
during forward trunk pertubation.” Exp Brain Res 148: (2003) 338-349.
Spitzer, W. O., Skovron, M. L., Salmi, L. R., Cassidy, J. D., Duranceau, J., Suissa, S. and Zeiss, E. “Scientific
monograph of the Quebec task force on whiplash-associated disorders: Redefining "whiplash" and its
management.” Spine 20(8S): (1995) 3S-73S.
Stemper, B. D., Yoganandan, N. and Pintar, F. A. “Gender dependent cervical spine segmental kinematics
during whiplash.” J Biomech 36: (2003) 1281-1289.
Stemper, B. D., Yoganandan, N. and Pintar, F. A. “Validation of a head-neck computer model for whiplash
simulation.” Med Biol Eng Comput 42: (2004) 333-338.
IRCOBI Conference – Graz (Austria) September 2004 280
Svensson, M. Y., Aldman, B., Hansson, H. A., Lovsund, P., Seeman, T., Suneson, A. and Ortemgren, T.
Pressure effects in the spinal canal during whiplash extension motion. International Research Council
on the Biomechanics of Impact (IRCOBI), (1993) Eindhoven, Netherlands, 189-200.
Szabo, T. J. and Welcher, J. B. Human subject kinematics and electromyographic activity during low speed rear
impacts. 40th Stapp Car Crash Conference, (1996) Albuquerque, NM, 295-315.
Tennyson, S., Mital, N. K. and King, A. I. “Electromyographic signals of the spinal musculature during +Gz
impact acceleration.” Orthop Clin North Am 8(1): (1977) 97-119.
TNO Automotive, T. N. MADYMO Theory Manual, version 6.0.
van den Kroonenberg, A., Philippens, M., Cappon, H., Wismans, J., Hell, W. and Langwieder, K. Human head-
neck response during low-speed rear end impacts. 42nd Stapp Car Crash Conference, (1998) Tempe,
AZ, 207-221.
van der Horst, M. J. Human head neck response in frontal, lateral and rear end impact loading - modeling and
validation. Eindhoven, The Netherlands, Technical University of Eindhoven: (2002).
Vint, P. F., Mclean, S. P. and Harron, G. M. “Electromechanical delay in isometric actions initiated from
nonresting levels.” Med Sci Sports Exerc 33(6): (2001) 978-983.
Vos, E. J., Harlaar, J. and van Ingen Schenau, G. “Electromechanical delay during knee extensor contractions.”
Med Sci Sports Exerc 23(10): (1991) 1187-1193.
Winkelstein, B. A., Nightingale, R. W., Richardson, W. J. and Myers, B. S. “The cervical facet capsule and its
role in whiplash injury.” Spine 25(10): (2000) 1238-1246.
Winters, J. M. and Stark, L. “Estimated mechanical properties of synergistic muscles involved in movements of
a variety of human joints.” J Biomech 21: (1988) 1027-1041.
Yang, K. H., Begeman, P. C., Muser, M., Niederer, P. and Walz, F. On the role of cervical facet joints in rear
end impact neck injury mechanisms. Motor Vehicle Safety Design Innovations, (1997), Society of
Automotive Engineers, Inc., 127-129.
Yoganandan, N., Pintar, F. A. and al., e. “Whiplash injury determination with conventional spine imaging and
cryomicrotomy.” Spine 26(22): (2001) 2443-2448.
Yoganandan, N., Pintar, F. A., Butler, J., Reinartz, J., Sances Jr, A. and Larson, S. J. “Dynamic response of
human cervical spine ligaments.” Spine 14(10): (1989) 1102-1110.
Yoganandan, N., Pintar, F. A., Stemper, B. D., Cusick, J. F., Rao, R. D. and Gennarelli, T. A. Single rear impact
produces lower cervical spine soft tissue injuries. International Research Council on the Biomechanics
of Impact (IRCOBI), (2001) Isle of Man, UK, 201-211.
Yoganandan, N., Pintar, F. A., Stemper, B. D., Schlick, M. S., Philippens, M. and Wismans, J. “Biomechanics of
human occupants in simulated rear crashes: Documentation of neck injuries and comparison of injury
criteria.” Stapp Car Crash Journal 44: (2000) 189-204.
Zhou, S., Lawson, D., Morrison, W. and Fairweather, I. “Electromechanical delay in isometric muscle
contractions evoked by voluntary, reflex and electrical stimulation.” Eur J App Phys 70: (1995) 138-
145.
IRCOBI Conference – Graz (Austria) September 2004 281
APPENDIX A: VALIDATION OF THE HEAD-NECK COMPUTER MODEL
Overall head to T1 angulation: experimentally obtained validation corridors are shaded, computer
model response is provided.
-15
0
15
30
45
60
75
50 100 150
He
ad
to
T1
an
gle
(d
eg
)
Time (msec)-15
0
15
30
45
60
75
50 100 150
-15
0
15
30
45
60
75
50 100 150
He
ad
to
T1
an
gle
(d
eg
)
Time (msec)
Segmental angulation: experimentally obtained validation corridors are shaded, computer model
response is provided.
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal a
ng
le (
de
g)
Time (msec)
C2-C3
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal a
ng
le (
de
g)
Time (msec)
C2-C3
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal a
ng
le (
de
g)
Time (msec)
C3-C4
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal a
ng
le (
de
g)
Time (msec)
C3-C4
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal a
ngle
(d
eg
)
Time (msec)
C4-C5
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal a
ngle
(d
eg
)
Time (msec)
C4-C5
-5
0
5
10
15
20
25
50 100 150
Segm
en
tal a
ng
le (
de
g)
Time (msec)
C5-C6
-5
0
5
10
15
20
25
50 100 150
Segm
en
tal a
ng
le (
de
g)
Time (msec)
C5-C6
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal an
gle
(d
eg)
Time (msec)
C6-C7
-5
0
5
10
15
20
25
50 100 150
Se
gm
en
tal an
gle
(d
eg)
Time (msec)
C6-C7
IRCOBI Conference – Graz (Austria) September 2004 282
Facet joint motion: experimentally obtained validation corridors are shaded, computer model response
is provided.
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C4-C5, Posterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C4-C5, Posterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C4-C5, Anterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C4-C5, Anterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C5-C6, Posterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C5-C6, Posterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C5-C6, Anterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C5-C6, Anterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C6-C7, Posterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C6-C7, Posterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C6-C7, Anterior
-1
0
1
2
3
4
5
50 100 150
Fa
ce
t jo
int m
otio
n (
mm
)
Time (msec)
C6-C7, Anterior