Much research on human tolerance to acceleration has been conducted at the United
States Air Force Aeromedical Research Laboratory (AFAMRL) in Ohio. The most extensive
work dealt with accelerations causing compression along the spine. This is positive +Gz or
"Eyeballs Down" acceleration. Less work has been conducted on accelerations perpendicular
to chest and parallel to the shoulders. This research has formed the basis of the dynamic
response criteria accepted by IMO.
To determine the injury potential of an acceleration field, Brinkley and Shaffer
(1971) introduced the concept of the Dynamic Response (DR). The basis of this concept is
the supposition that each body axis can be idealized as an independent single
degree-of-freedom spring-mass system that is subjected to known seat accelerations
(Brinkley and Shaffer, 1984). This model is shown in Figure 6.4. It was originally
developed to evaluate the effects of acceleration along the spine, but has been expanded to
evaluate the effects of acceleration perpendicular to the chest and parallel to the shoulders.
The dynamic response is computed by:
DR( t ) = - (ω2
x t) / g (6.3)
where omega is the undamped natural frequency for the axis being studied, t) is the
displacement time-history of the body mass relative to the seat support, and g is
gravitational acceleration. Values for the
natural frequencies in each co-ordinate axis have been found through research conducted at
AFAMRL. These values for a 50th percentile 28 year old male in a fully restrained seat
and harness are presented in
Figure 6.4; Independent single degree-of-freedom representation of the human
The DR computed with Equation 6.3 represents an equivalent static acceleration of
the body mass in an undamped system. The peak value of the DR curve is called the
Dynamic Response Index (DRI). It can be used as an indicator of the potential for
acceleration forces to cause human injury.
Table 6.1: Parameters of the dynamic response model
|Coordinate Axis||Natural Frequency (rad/s)||Damping Ratio|
Brinkley and Shaffer (1984) defined three risk levels for acceleration forces directed
along the spine. These risk levels are characterized as high, moderate, and low. They
relate to a 50%, 5% and 0.5% probability of injury, respectively. The 50% probability of
spinal injury is the highest rate observed for USAF ejection seats. It should be noted that
there was no spinal chord damage associated with these injuries. The moderate risk level,
which is currently used in USAF ejection seat design, is midway between the high risk
and low risk levels. The low risk level corresponds to acceleration conditions used
routinely without incident during test conducted with volunteers at the AFAMRL. The three
injury curves presented by Brinkley (1985) for the +z axis are shown in Figure 6.5. Each
curve, which represents a constant DRI at the appropriate risk level, was computed from
half-sine acceleration impulses acting at the seat support. The DRI limits presented by
Brinkley (1985) for each risk level are shown in Table 6.2.
Figure 6.5 Three risk Level for Acceleration Acting in the + Z axis (Brinkley 1985).
The risk levels for the +/-x, +/-y, and -z axes were determined without the benefit
of a statistically based method such as that used for the +Z axis (Brinkley and Shaffer,
1984). The high risk levels were determined by calculating the peak response of the
mathematical model to acceleration conditions known to cause major injuries or potentially
serious sequelae. he low risk levels were determined on the basis of calculated model
responses to acceleration conditions that have been used numerous times for noninjurious
tests with human subjects in research laboratories. The moderate injury level was assigned
as the midpoint between the high and low levels. The DRI limits for these axes are
presented in Table 6.2.
Table 6.2: DRI Limits for three risk levels
|Coordinate Axis||DRI limits (G's)|
The response limits presented in Table 6.2 are for single axis accelerations.
Normally, components of acceleration are acting in each co-ordinate direction
simultaneously. The effects of multi-axial acceleration can be evaluated with an ellipsoidal
envelope. The boundaries of the envelope in each direction are the DRI limits presented
in Table 6.2. The seat accelerations to which the occupant is subjected, then, are limited
DRIx, DRIy, and DRIz in Equation 6.4 are the limiting DRI's in the x, y, and z
co-ordinate directions, respectively, for a particular risk level.
DRx, DRy, and DRz are the calculated dynamic responses of the body mass in the same
The number of computations required to evaluate an acceleration field can be
reduced if the apparent relative displacement of the body mass with respect to the seat
support is used directly to determine acceptability instead of computing the DR. This
approach is valid because, as shown in Equation 6.3, the DR is a linear function of the
The displacements permitted by the IMO criteria are presented in Table 6.3. These
allowable displacements were computed from the DRI limits presented in Table 6.2 and the
natural frequencies presented in Table 6.1. The "Training Condition" corresponds to a
0.5% probability of injury. This was deemed to be the maximum acceptable for training
exercises because these acceleration forces will be experienced several times. The
"Emergency Condition" corresponds of a 5% probability of injury. This level was deemed
acceptable in potentially life threatening situations.
Table 6.3: Suggested displacement limits for lifeboats
When relative displacements are used as the basis for determining the acceptability of
an acceleration field, the effects of multi-axis accelerations can be evaluated with the
Combined Dynamic Response Ratio (CDRR). The CDRR represents a displacement
envelope that is analogous to the acceleration envelope implied in Equation 6.4. It is
Sx, Sy, and Sz in Equation 6.5 are the allowable displacements in the x, y, and z
co-ordinate directions, respectively, for a particular risk level. are the apparent computed
relative displacements of the body mass with respect to the seat support. The peak value
of the CDRR time-history is called the CDRR Index. A CDRR Index that is less than or
equal to unity indicates the particular risk level has not been exceeded. When this analysis
is performed, the acceleration data are not filtered. If an acceleration force does not have
a significant influence on the human, it should not have a significant influence on the
response of the model.
Although Brinkley has shown a good correlation between spinal injury and DRI,
experience with Royal Air Force ejections indicates that the incidence of spinal injury is
not as well correlated with the DRI. Anton (1991) has presented as series of 223 "within
envelope" ejections in which the discrepancy between predicted injury rate and observed
injury rate is very wide. The DRI indicated that the injury rate should be about 4% but
the actual injury rate was 30-50% depending on the type of seat used.
A possible explanation for this discrepancy is the type of harness used in military
aircraft in the United Kingdom. The preferred harness type is the simplified combined
harness (SCH) which provides restraint in normal circumstances but becomes the parachute
harness after ejection. The harness is part of the seat equipment. The torso harness, on
the other hand, is favoured by the US forces. It is fitted to the crew member and worn
out to the aircraft where it is attached to the seat.
The torso harness is recognized to provide slightly better coupling between the
man and the seat. This coupling is important in reducing "dynamic overshoot," the effect
in which the occupant experiences greater dynamic forces than those applied to the seat.
There is, in essence a second mass spring damper system operating at the interface
between the seat and occupant. This second spring damper can, however, be incorporated
into the dynamic response model but knowledge of seat coupling and cushion stiffness is
The effects of seat/occupant coupling, and the apparent effect of this coupling on
observed injury rates, indicates a strong need for free-fall lifeboats to be equipped with a
properly designed harness system. The issue of harnesses is discussed later in this paper.