From an ongoing comprehensive reading of published documents, a format for the presentation of fully researched results has emerged in the form of four primary categories:
1. Common types of injuries associated with undue mechanical force on aircrew members in DoD aviation operations
2. Their sites and frequency of occurrence, with mention of underlying causes of injury, where known
3. Estimated tolerances of aircrew members beyond which such injuries (major, fatal) are most likely to occur
4. Estimated DoD operational, equipment and health care costs associated with these injuries with suggestions of reasonable countermeasures and/or improvements to current systems of protection compatible with DoD mission requirements, and projected cost savings if such remedial proposals are successfully implemented.
At this point, it is proposed to set out below in very succinct form a brief summary of the results of the current preliminary literature survey with regard to the first three categories above.
1. Fracture injury types – mainly anterior-lip crush fractures due to hyperflexion
a) Upper lumbar and lower thoracic spine: T11 to L2 most common sites account for about 20% of all ejections (7% – 47% for other data bases).
b) Cervical fractures: C2, C5 and C6 most common sites. These fractures are 5 to 7 times less common than a) but are sometimes fatal. They are not continuously proportional to applied mechanical force. For example, a relatively wide range of forces on C1/C2 may be easily tolerated until and unless separation of the axis/atlas joint occurs, which can be instantly fatal. Cervical fractures comprise approximately 2% of ejections from both jettisoned and through-the-canopy ejection escape systems.
c) Head/neck injuries: are mainly fractures and dislocations in 70%-77% of all ejection fatalities, 3/4 of which are categorized as “multiple trauma”. Such fatal injuries are strongly associated with through-the-canopy ejections. However, for ejections with pre-jettisoned canopies, more deaths may then occur post-egress from striking of aircraft parts. There appears no single causal factor for ejectees, but additional injuries relate to system malfunction before and during parachute opening and seat striking head after seat separation, entanglement with seat before parachute opening. There is also some evidence of “whip-lash” type injuries associated with a rapid onset of G-forces, during which there may be uncontrolled extension, then flexion of the head/neck complex. This may occur in under 50 ms, before the neck muscles have had time to activate to offer any resistance (normal muscle reaction time is between 5 and 20 ms). The latter is serious, as it is an indicator of vestibular system pathology.
2. Injury statistics – Fatality rates vary widely: 5.9% – 53.7% (below 500 ft.)
a) Overall fatality rate for ejections from fighter and attack aircraft:
< 10% for modern escape systems,
= 10.7% for through-the-canopy systems without fragmentation cords
It is noted that even an incremental fatality rate of 1% should be considered unacceptable if it can be remedied simply (eg. explosive cord to fragment canopy prior to ejection through it). During the 1950 – 1960’s, less than 1/3 of ejections took place during highly dynamic uncontrolled flight. However, during the 1980’s, with the advent of more modern and electrically controlled aircraft, that figure rose to approximately 2/3 (Guill et al. 1989). In his U.S. Air Force injury summary for FY 1989, Campbell (1990) identifies the leading causes of fatalities in aircraft equipped with escape systems to be out-of the-envelope ejection and collision with the ground.
b) Thoraco-lumbar fracture statistics:
Visuri & Aho (1992) – amongst 19% of survivors of ejections
Sandstedt (1989) – amongst 18-27% depending upon posture and flight conditions. Alignment of head, neck, spine reduces fractures,
McCarthy (1988) – 21% of survivors in take-off/landing ejections and same for ejections above 500 ft. altitude, but 2.5 times higher for altitudes below 500 ft.
Navy Safety Center, vol. IV (1969-1979) – 28% for through-the-canopy Navy ejections – 7% for jettisoned canopies overall
c) Spinal fracture statistics by NATO country:
France: Auffret & Delahaye (1975): 10%- 47% of surviving ejectees, of which 37% with fractures at T12 or L1 vertebrae,
Italy: Rotondo (1975): 36% of ejectees, nearly uniformly distributed over T7 to L4, except fracture frequency 4 times greater for T12 and L1,
Nutall (1971): T11 to L2 fractures most likely – (1957) study: 2.2%, and
(1965) study: 3.8% spinal fractures of total ejections
Greece: Symeonides (1971): 18% of ejectees, fractures in T11 to L2 region,
Other: Henzel (1967): T10 to L1 are most common fracture sites,
Jones (1964): T12 and L1 are the most common fracture sites.
d) Windblast and parachute opening shock injuries:
Windblast injuries: are mostly due to limb flail.
Parachute opening: large transient Gz forces, depending upon constraints
There is no evidence to support “ejection-associated” neck injuries primarily produced by parachute opening shock (Guill et al. 1989).
Parachute landing falls: appear to account for the largest single component of fatalities for ejections (92/833) during the years 1978-1992.
In marginal or gusty conditions, the oscillation of a round canopy may inhibit control and safe landing, often at descent rates of 700 ft/min with 250 lb at sea level. Ram-air parachute canopies offer an interesting alternative with docile flight characteristics, according to Precision Aerodynamics, Inc. of Dunlap, TN.
Studies by Brinkley & Shaffer (1971) indicate that ejection boost acceleration is the primary cause of major injuries in 84% of F-4 ejections, and 12% of post-ejection injuries are due to limb flailing, while only 5% occur from parachute opening shock (corresponding to 1.3% of 384 ejections).
e) Helmet impact with canopy:
Brinkley et al. (1975) report on 1303 ground impacts during the period 1963-1967 for jet fighters (65.5% survival rate) and helicopters (93.4% survival rate). They found fatal head injuries were highest in helicopters (6.1%) while major head injuries were highest in fighter aircraft (13.6%). The rate of loss of helmets after ejection increased 2 1/2-fold if the visor was left open during ejection. Cervical ejection injuries were observed to decrease by 14.6% without the helmet, possibly due to the absence of added head weight.
From the overall findings, however, it would appear that the protection provided by harnesses and lap-belts outweighs the protection of helmets. In other words, it seems that the head is not vulnerable to injury during ejection as long as it does not strike a surface. However, during ejection, rapid tumbling of the seat during tip-off can cause high speed separation with collision of the helmet with the head rest (Guill et al. 1989).
3. Documented causes of types of injury
1. Spinal compression injuries have recently been studied by Shirazi-Adi et al. (1996) in terms of the natural exertion of stabilizing moments at the upper end of the spine and of pelvic rotation at the lower spine. On the basis of their nonlinear 3-D finite element code predictions, they postulate the existence of an optimal combined set of end conditions which stabilizes the passive thoraco-lumbar spine, thus permitting it to support much larger compression loads with minimal displacements. It is suggested that by suitable positioning of the center of gravity of the upper body in the presence of pelvic rotation influencing the spatial postion of T1, one can effectively regulate this optimal set of end conditions in such a way as to enhance the load-bearing capacity of the spine in compression. The role of the central nervous system in maintaining this postural control is not yet fully understood. However, the seat and seat-back of the aircrew members may conceivably be fitted to assist in this regulation of posture prior to emergency ejection.
2. Head-neck-spinal injuries due to unsatisfactory pre-escape position of aircrew member: proper alignment of head, neck and spine is critical during ejection if injury is to avoided. De-positioning may often occur because of very high Gz forces on the body during aircraft maneuvering. Automatic restraint systems actuated prior to ejection can cause mid-thoracic and cervical fractures. One should maintain a 135° angle between torso and thigh for proper alignment of thoracic vertebrae (Auffret et al. 1975) during the application of ejection boost forces. The harness must hold its position under high G maneuvers.
Nightingale et al. (1996) have reported the results of a truly outstanding experimental study into the time sequence of events, measured in milliseconds, which lead to injury of the cervical spine following axial impact of the head onto either rigid or padded surfaces. In their experiments, they dropped eleven whole unembalmed human head-cervical spine complexes to produce impact velocities of the order of 3.2 m/s. The effective added torso mass contributing to the momentum of the head upon impact was estimated at 16 kg. During phase 1 lasting about 4.3 ms, it was observed that the falling head, which was neutrally aligned with the cervical spine, stopped in about 4 ms. after impact. The head forces of impact (significantly lower for padded impacts than for rigid surface impacts) were augmented by the inertial loading of the torso, with no corresponding neck force yet developing. As the head rebounded from the impacting surface, the T1 vertebra was compressed, sustaining injury near the end of phase 1, after about 8 ms. In the subsequent phase 2 lasting about 27 ms, during which the inertial follow-through of the torso continued to load both the head and the cervical spine. At the first instant of injury to the cervical spine T1/T2, the load on the neck can exceed the instantaneous load on the head which is relieved during the rebound phase. No significant difference in neck loading was noted between padded and rigid impacts. At 3 to 8 ms after head impact, the cervical spine was observed to buckle dynamically, both in a first-order but also higher order mode (sometimes with rapid “snap-through” from one equilibrium buckling configuration to another), as evidenced by an abrupt decrease in measured compressive loading and increased bending moment and transverse deflection. The bending stresses so developed in the cervical spine are likely responsible for the crushing anterior lip fractures observed and which correspond to the same sites noted in ejection injuries. These appear to be extension-type injuries in C2-C3-C4-C5 and the connecting ligaments and flexion-type lesions in C6-C7-T1 and their connecting ligaments. These experiments reinforce the importance of end conditions provided by the head and torso orientations on the critical loading of the cervical spine. Although absent from these cadaver experiments, the active role of the neck muscles may be to absorb and stabilize energy during trauma, particularly during the extension of the cervical spine. However, in previous studies reported by Foust et al. (1971) and Schneider et al. (1975), it has been noted that the muscle reflex times of 50 to 65 ms are far too slow to influence 20 ms time-frame during which injury is inflicted on the cervical spine. Due to its slenderness, the cervical spine will buckle before compression failure occurs as a result of head impact.
3. Pre-existing back problems of ejecting aircrew members may lower their tolerance to spinal fractures. The latter do not appear to be particularly sensitive to the anthropometry of the ejectees, however.
4. Cervical column “whiplash” (extension-flexion), occasioned by in-flight maneuver forces, is potentially injurious. Rapid inertial loading induced by whipping motion of the head and neck during the rapid onset or cessation of high G, may occur during combat maneuvers. It is thought that the flexion phase, following initial extension of the neck during cervical whiplash injury, may be far more damaging than the extension phase. Flexion probably causes far more persistent damage to the central nervous system by pulling on the cervical medulla, that force being conducted up to the medulla oblongata and to the brain stem. This mechanical action may lengthen these structures by up to 5 cm, leading to extensive damage of the central nervous system. Although only 28% may suffer any immediate limitations in head-neck movements, up to 94% have cervico-brachialgia, 88% have radiating headaches, 79% vertigo and dizziness (over-excitation of the cervical proprioceptors) and 77% show evidence of cervical induced nystagmus (cf. Oosterveld et al. 1989). However, it remains to be seen whether the severity of whiplash injuries reported by Oosterveld for automobile rear-end collisions obtains in non combat or combat aviation operations. First indications would appear to negate the development of such acute degrees of extension of the brainstem structures in aviation operations.
Nystagmus, an involuntary rapid movement of the eyeball (horizontal, vertical, rotary or mixed) is a clear sign of vestibular system pathology. Both spontaneous and positional (> 5°/sec), it is found in about 60% of automobile whiplash cases, whereas its normal incidence in the general population is less than 2%. Some 80% of victims of whiplash injury have bilateral gaze, indicating brainstem pathology, while 82% show signs of visual pursuit disturbances, signaling dysfunction of the central nervous system with cerebellar pathology.
Long after the occurrence of whiplash injury, cervical soft tissue lesions and ruptures of muscles and ligaments in the neck may change morphologically to inflammatory granulation tissue with scarring and degeneration of nerves in the cervical area. Restrictive as it may appear, the remedy may well be to design innovative head restraints to surround the head of the aircrew members, rather than serving simply as a back head rest (which effectively only protects against over-extension).
4. Estimated tolerances of aircrew members
Before embarking on this study, it should be pointed out that the task of relating levels of injuries actually sustained in aviation operations to magnitudes of loading forces is not straightforward. Attempts have been made to define “injury criteria” in terms of various levels of microscopic and macroscopic idealizations. For example, these include:
1. stress/strain in tissues as a function of localized tissue injury
2. forces, moments and accelerations applied to body segments to injury of that segment
3. dynamic loading histories to whole body injury
Type 1 appears to be impractical to realize because of the absence of:
a) sufficiently detailed accident data for the cellular tissue level
b) suitably realistic manikins available to test such detail experimentally.
For example, more detailed tests are needed to measure compressive anterior/posterior lip loads at T12 and L1 vertebrae in manikin simulations. Consistent and statistically significant compressive strength data are needed from 20 to 30 year-old cadaveric specimens subjected to typical aircrew-encountered loadings, especially for the cervical spine at levels C1 to C5, or the establishment of valid scaling between successive Cn vertebral levels. However, many published papers contain estimates of ultimate compressive strength of isolated vertebral bodies for both fast (0.9 m/s) and slow (9´10-5 m/s) loading rates.
Accordingly, the only remaining correlations which are practically utilizable are types 2 and 3 above. However, pre-1980 data on whole body ejection dynamics may lack sufficient detail for the re-design of current ejection systems.
Generally accepted acceleration tolerances to avoid spinal fractures during ejection for aircrew members who are properly aligned with the acceleration vectors are 20~25 G.
There is much less agreement on the tolerable duration and end-of-stroke velocity during ejection. The accepted ranges of: maximum rate of onset: 200 ~ 500 G/s
minimum pulse duration: 100 ~ 230 ms
maximum change in velocity: 20 ~ 60 fps
Mertz and Patrick (1972) give head-neck response envelopes for flexion and for extension at various tolerance levels. These are graphed as moment at occipital condyles vs. head rotation or position. It has been suggested that voluntary tolerance levels are about one-half to two-thirds of injury tolerance levels. For example, weightlifters will not voluntarily execute a lift producing lumbar compressive forces greater than two-thirds of their ultimate lumbar compressive strength. However, it should be noted that differences in postures and the dynamics of the loading history also play an important role here.
Tolerance to neck damage is not merely a function of mechanical loading alone. Neurological damage and severe strains can occur even without having attained loading levels which fracture ligaments or cerebral vertebrae. The latter voluntary and cadaveric tolerance levels have been reported by Mertz and Patrick (1972) and by Patrick (1987) for both dynamic moments and static force applications at the occipital condyles. However, no reports have yet been identified which relate specifically to ejection-generated cervical fractures. Progressive cervical degeneration (osteoarthritis) sustained by aircrew members following prolonged and repeated exposures to sustained high G accelerations (Gillen et al. 1989) may initially be excluded from the scope of the present investigation, although its effects on aircrew member performance in tracking enemy targets in-flight may be significant.
Pintar et al. (1989) have also reported upper cervical injury tolerances in compression-extension and lower cervical injury tolerances in compression-flexion by direct testing of human cadaveric head-neck complexes. They find failure loads of 305 ~ 812 lb. and strains at failure of 0.04 ~ 0.26.
Melvin (1977) has reported cervical vertebral fracture-producing axial compressive loads greater than 1000 lb; that is, much greater than the voluntary tolerance of 250 lb.
McElhaney et al. (1983) estimate that a dynamic compressive load of about 1400 ~ 1800 lb. is required to fracture C5.
The lowest load required to crush vertebral body C5 was reported in the SAE Info Report J885 JUL86 to be 1620 lb., while loads of 1990 lb will crush the discs between C3-4, C4-5 and C5-6, two transverse processes and the T2 vertebral body. Again, fractures can occur at even one-half these loads if the head, neck and torso are not aligned parallel with the ejection acceleration vector.
Shear force injuries are most likely between the occipital condyles and C2. These shear forces need to be measured in the upper neck of realistically designed manikins. However, injury due to dynamic impact loading of the neck will depend on many variables, such as: impact location, line of action of the forces, concentrated vs. distributed loading, initial neck curvature, energy levels and the presence of protective gear.
Ejection forces and accelerations exerted during a transonic 720 KEAS test have been measured on a fully instrumented manikin, and its ejection seat as a function of time (Frisch et al. 1989). During the 15 Gz catapault phase, lasting 150 ms to 275 ms, a lumbar load of 725 lbs was measured at the pelvic-flexible lumbar spine interface. Compressive loads of +85 lbs developed at the neck-occipital joint interface during this phase, but reverse to tensile forces of -500 lbs at 300 ms due to the windstream coming up the manikin chest cavity (backward reclining seat) and interacting with the chin. The head is violently rotated backward during this aerodynamic interaction, forcing it against the headbox, which arrests any further rotation of the head. Further airstream forces exerted under the chin puts the cervical spine into tension, independently of any subsequent changes in aerodynamic lift properties of the head and neck system.
The best current estimate of axial compression tolerance for adults for various injury levels may nonetheless be presented as:
Serious injury: forces > 250 lb. for a duration of Åt > 30 ms
> 850 – 20 Åt lb. for Åt < 30 ms
which corresponds to an AIS level 5 (critical) injury. As noted in section 1 b) above, there is no continuous scaling of level of injury with mechanical force. For example, no injury might occur as long as C1 and C2 remain together, whereas their sudden separation could prove fatal.
“Pocketing” of the head/neck upon head impact can increase the potential for injury. Accordingly, it has been suggested that neck injuries may be diminished by reducing the friction between helmet and struck surface, especially for crown impacts.
The loading tolerance to dislocate the lower cervical spine bilaterally is much lower than the axial load tolerance for vertebral compression fractures. Huelke et al. (1985) believe that many flexion-type injuries of the neck (head bowing to the chest) occur before gross head motion starts. It would therefore appear useful to estimate the time-frame for such injuries to occur as a means of establishing a realistic time-sequence on which to base interpretations of filmed events with manikins or cadavers, for example.
McElhaney et al. (1989) have undertaken a comprehensive experimental investigation of the flexion, extension and lateral bending of the cervical spine. They found that bending is significantly influenced by: the direction of the bending moment, the type of end restraints on the cervical spine and the corresponding shear forces produced there, the magnitude of the deformation and its previous history, and the eccentricity of the spinal loading. It was concluded that bending is the primary deformation mode. Axial loading is a poor indicator of failure stresses in the cervical spine. Damage may be best visualized by MR and this may be a feasible non-invasive tool for in vivo use on aircrew members.
The following tolerances have been reported by McElhaney et al. (1989) for the cervical spine:
Compression breaking loads 1.47 – 2.16 kN (Messerer 1880)
Quasi-static compression failures 0.645 kN (Sances et al. 1981)
Dynamic compression-flexion failures 1.78-4.45 kN (Sances et al. 1981)
Forward dislocations 1.32-1.42 kN (Bauze et al. 1978)
Time-varying compressive loads 1.93-6.84 kN (McElhaney et al. 1989)
Failure loads 3.2-10.8 kN (Nusholtz et al. 1981)
Small eccentricities in the load axis relative to the cervical spinal axis can alter the buckling load from posterior to anterior. Roaf (1960) believes ligamentous rupture results only from rotation and/or shear forces, and not from compression, flexion or extension. Hodgson et al. (1981) measured local strains of cervical spine during head impact and found that off-axis torsional and transverse shear were important for axial response. An extensive bibliographic review has been provided by Sances et al. (1981).
Both Seemann et al. (1986) and McElhaney (1989) conclude that although there may be satisfactory agreement in the dynamic response of the Hybrid III neck with the human neck for some bending modes, a fundamental difference between the two appears in terms of changes in stiffness with displacement rate and hysteresis. However, Seemann et al. (1986) suggested that improved agreement in the -Gx simulations may be possible by relocating the neck/torso joint of the Hybrid III model.
Tarrière et al. (1989) report on the statistics and tolerances of both men and women to cervical lesions produced by head impact in automobile collisions. Frontal, lateral and rear impacts are considered. No cervical injuries were reported in the absence of head impact. Neck injuries linked with hyperextension and of severity AIS >1 are two to three times for probable in women than in men. The authors suggest these dramatic differences between the sexes are due to differences in the cervical morphology between men and women and the differences in positioning of the two sexes relative to their respective restraint systems. It was found that 70% of the cervical lesions were situated in the brain stem. The Head Injury Criterion (HIC) was shown not to be a reliable predictor of head injury in these tests, as the mean HIC and even the ranges of HIC were very similar for both minor and major injuries in frontal and lateral impacts. The authors suggest that what is needed is a combination of several parameters, such as angular velocity and angular acceleration of the head, to be incorporated into an improved injury criterion. From their research with volunteer boxers during training fights, the upper bounds for tolerance for no physical or physiological effects to the head appear to be: a) angular velocity < 48 rad/s, and b) angular acceleration < 16,200 rad/s2 for movements of the head.
Human tolerances for impact to the lower extremities were compiled by Snyder (1971) for the: lower leg (tibia) @ 1000 ~ 3000 + lb (mean @ 1500 lb.)
knee (patella) > 1050 ~ 3850 lb load
upper leg (femur) @ 20 ~ 50 ft-lb energy
Flail injuries have been reported by Payne et al. (1974). It is believed that flail injuries need never occur in a properly designed ejection seat at any point in the operational envelope of today’s aircraft. However the probability of survival curves presented may not be relevant to modern aircraft. Buschman et al. report on USAF non combat operational experience for the period of 1957-1970, for which they estimate 13% flail and possible flail injuries, representing 7% of all ejections at median airspeeds of 225 knots.
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Expert 128475 is a former university professor of biomedical and human factors engineering, former professor of medicine, and former professor of mechanical and aerospace engineering. His expertise includes analyses of vehicular, industrial, recreational, commercial, and sports-related injuries in addition to the evaluation of protective devices and safety systems such as seat belts, air bags, clothing, guards, and helmets.