Effects Of Pelvis Impact Angle And Hip Muscle Forces On Hip Fracture Risk During A Fall Using An Advanced Hip Impact Simulator

Over 90% of hip fractures in older adults are caused by falls [1]. Whether a given fall will cause hip fracture depends on bone strength, and on the impact force and stress applied to the bone during impact [2]. Improved understanding is required on how peak bone stresses during a fall depend on the mechanics of a fall, and on the state of contraction at the moment of impact of the muscles spanning the hip. Recently, Choi et al showed that, for lateral impact to the hip, peak stresses decrease with increases in hip abductor muscle force [3]. In the current study, we used an advanced hip impact simulator to examine the independent and interacting effects of both hip muscle force and pelvis impact angle on peak bone stresses during a fall

Influence Of Pelvis Impact Angle During A Fall: On The Protective Benefit Of Hip Protectors

Over 90% of hip fractures are due to falls [1]. Laboratory measures have shown that wearable hip protectors reduce impact forces to the proximal femur during a simulated sideways fall on the hip [2, 3]. However, clinical evidence suggests that hip fractures still occur when hip protectors are worn [4]. Furthermore, while falls in real life result in a variety of impact configurations, biomechanical tests to date have focused only on lateral impact to the pelvis. In the current study, we examined how the force reduction provided by wearable hip protectors is affected by pelvis impact configuration during simulated sideways falls

Pressure Distribution Over the Palm Region During Forward Falls on the Outstretched Hands

Falls on the outstretched hands are the cause of over 90% of wrist fractures, yet little is known about bone loading during this event. We tested how the magnitude and distribution of pressure over the palm region during a forward fall is affected by foam padding (simulating a glove) and arm configuration, and by the faller’s body mass index (BMI) and thickness of soft tissues over the palm region. Thirteen young women with high (n=7) or low (n=6) BMI participated in a “torso release experiment” that simulated falling on both outstretched hands with the arm inclined either at 20° or 40° from the vertical. Trials were acquired with and without a 5 mm thick foam pad secured to the palm. Outcome variables were the magnitude and location of peak pressure (d, θ) with respect to the scaphoid, total impact force, and integrated force applied to three concentric areas, including “danger zone” of 2.5 cm radius centered at the scaphoid. Soft tissue thickness over the palm was measured by ultrasound. The 5 mm foam pad reduced peak pressure, and peak force to the danger zone, by 83% and 13%, respectively. Peak pressure was 77% higher in high BMI when compared with low BMI participants. Soft tissue thickness over the palm correlated positively with distance (d) (R=0.79, p=0.001) and force applied outside the danger zone (R=0.76, p=0.002), but did not correlate with BMI (R=0.43, p=0.14). The location of peak pressure was shunted 4 mm further from the scaphoid at 20° than that of 40° falls (d=25 mm (SD 8), θ= −9° (SD 17) in the 20° falls versus d=21 mm (SD 8), θ= −5° (SD 24) in the 40° falls). Peak force to the entire palm was 11% greater in 20° compared with 40° falls. These results indicate that even a 5 mm thick foam layer protects against wrist injury, by attenuating peak pressure over the palm during forward falls. Increased soft tissue thickness shunts force away from the scaphoid. However, soft tissue thickness is not predicted by BMI, and peak pressures are greater in high individuals than that of low BMI individuals. These results contribute to our understanding of the mechanics and prevention of wrist and hand injuries during falls

Pressure Distribution Over the Palm During Falls on the Outstretched Hands

Over 90% of wrist fractures are caused by falls on the outstretched hands (Palvanen et al., Osteoporos Int, 2000). Along with bone strength, fracture risk depends on the magnitude and distribution of force to soft and hard tissues during impact. In the current study, we examined how pressure distribution over the palm during a fall is affected by impact configuration, body mass index (BMI), palmer soft tissue thickness, and a 5 mm thick foam pad (simulating a protective glove)

Development and Validation of a Questionnaire for Analyzing Real-Life Falls in Long-Term Care Captured On Video

Background Falls are the number one cause of injuries in older adults, and are particularly common in long-term care (LTC). Lack of objective evidence on the mechanisms of falls in this setting is a major barrier to prevention. Video capture of real-life falls can help to address this barrier, if valid tools are available for data analysis. To address this need, we developed a 24-item fall video analysis questionnaire (FVAQ) to probe key biomechanical, behavioural, situational, and environmental aspects of the initiation, descent, and impact stages of falls. We then tested the reliability of this tool using video footage of falls collected in LTC. Methods Over three years, we video-captured 221 falls experienced by 130 individuals in common areas (e.g., dining rooms, hallways, and lounges) of two LTC facilities. The FVAQ was developed through literature review and an iterative process to ensure our responses captured the most common behaviours observed in preliminary review of fall videos. Inter-rater reliability was assessed by comparing responses from two teams, each having three members, who reviewed 15 randomly-selected videos. Intra-rater reliability was measured by comparing responses from one team at baseline and 12 months later. Results In 17 of the 24 questions, the percentage of inter- and intra-rater agreement was over 80% and the Cohen\u27s Kappa was greater than 0.60, reflecting good reliability. These included questions on the cause of imbalance, activity at the time of the fall, fall direction, stepping responses, and impact to specific body sites. Poorer agreement was observed for footwear, contribution of clutter, reach-to-grasp responses, and perceived site of injury risk. Conclusions Our results provide strong evidence of the reliability of the FVAQ for classifying biomechanical, behavioural, situational, and environmental aspects of falls captured on video in common areas in LTC. Application of this tool should reveal new and important strategies for the prevention and treatment of falls and fall-related injuries in this setting

Performance of a Hip Protector Depends on its Position During a Fall

Hip protectors are designed to attenuate and redistribute the force applied to the hip region during a fall, and thereby reduce risk for hip fracture [1]. However, little information exists on the effectiveness of hip protectors in achieving these goals, and how this is altered by displacement of the hip protector relative to the greater trochanter (GT). In the current study, we tested these issues

A Principal Components Analysis Approach to Quantifying Foot Clearance and Foot Clearance Variability

Low foot clearance and high variability may be related to falls risk. Foot clearance is often defined as the local minimum in toe height during swing; however, not all strides have this local minimum. The primary purpose of this study was to identify a nondiscrete measure of foot clearance during all strides, and compare discrete and nondiscrete measures in ability to rank individuals on foot clearance and variability. Thirty-five participants (young adults [n = 10], older fallers [n = 10], older nonfallers [n = 10], and stroke survivors [n = 5]) walked overground while lower extremity 3D kinematics were recorded. Principal components analysis (PCA) of the toe height waveform yielded representation of toe height when it was closest to the ground. Spearman\u27s rank order correlation assessed the association of foot clearance and variability between PCA and discrete variables, including the local minimum. PCA had significant (P \u3c.05) moderate or strong associations with discrete measures of foot clearance and variability. An approximation of the discrete local minimum had a weak association with PCA and other discrete measures of foot clearance. A PCA approach to quantifying foot clearance can be used to identify the behavioral components of toe height when it is closest to the ground, even for strides without a local minimum

An analysis of the effect of lower extremity strength on impact severity during a backward fall

to 90 J of impact energy into horizontal (as opposed to vertical) kinetic energy. Declines in joint strength reduced the effectiveness of mechanisms (1) and (3), and thereby increased impact severity. However, even with reductions of 80 percent in available torques, KE v was attenuated by 50 percent. This indicates the importance of both technique and strength in reducing impact severity. These results provide motivation for attempts to reduce elderly individuals' risk for fall-related injury through the combination of instruction in safe falling Introduction Among elderly individuals living in the United States, falls are the number one cause of nonfatal injury, and number two cause of injury-related death ͓1͔. They account for approximately 90 percent of hip and wrist fractures in this age group ͓2,3͔, and nearly 40 percent of traumatic vertebral fractures ͓4͔. The annual medical costs in the U.S. associated with hip fractures alone is approximately $8.9 billion ͓5͔. While one's risk for fall-related fracture is influenced by bone density and fall frequency, growing epidemiological evidence suggests that it depends most strongly on the mechanics of the fall. This makes sense from a biomechanical perspective, given that the energy and force capable of being generated during a fall from standing height substantially exceed values required to fracture the proximal femur or distal radius ͓6-10͔. The most important determinant of injury risk during a fall is impact location. Direct impact between the hip and the ground increases elderly individuals' risk for hip fracture by approximately 30-fold ͓11-13͔. Impacting the outstretched hand reduces hip fracture risk by about threefold and increases risk for wrist fracture by approximately 20-fold ͓11͔. Fall direction influences impact configuration and injury risk, with sideways falls creating the highest risk for hip fracture, and backward falls creating the highest risk for wrist and vertebral fracture ͓4,11,14͔. Biomechanical considerations suggest that injury risk also associates with impact velocity. For example, our previous ''pelvisrelease'' experiments ͓15,16͔ indicate that, while body mass and soft tissue thickness have important influences, impact velocity is by far the strongest determinant of femoral impact force during a fall on the hip. While it has been impossible to include this variable directly in epidemiological studies of fracture risk, previous studies have shown associations between hip fracture and both fall height ͓2,12͔ and body height ͓17͔, presumably due to the effect of these variables on impact velocity. Finally, declines in lower extremity strength substantially increase hip fracture risk in the event of a fall ͓13,18 -20͔. While the mechanism underlying this association is unknown, some authors have suggested that it relates to one's ability to absorb energy in the lower extremity muscles during the descent phase of the fall, and thereby reduce the kinetic energy and velocity of the body at impact ͓21͔. Consideration of acts such as sitting and squatting indicate the potential effectiveness of this mechanism. When a typical adult descends from standing to sitting, the total potential energy of the body decreases by approximately 200 J. However, due to energy absorption in the lower extremity muscles during descent, the body's downward velocity and kinetic energy at the instant of chair contact tend to be minimal ͓22͔. Õ Vol. 123, DECEMBER 2001 Copyright © 2001 by ASME Transactions of the ASME chanical studies suggest that young subjects utilize this mechanism to reduce impact severity during falling, since observed contact velocities are well below those predicted by simple mathematical models of free fall ͓22-25͔. However, little is known regarding the falling techniques that minimize impact severity, and how the efficacy of these depends on lower extremity strength. Accordingly, the main goal of the present study was to examine the theoretical effect on fall severity of lower extremity muscle contractions during descent. A related goal was to determine how one's ability to reduce impact severity is affected by age-related declines ͑or exercise-induced enhancements͒ in lower extremity strength. Given the complex out-of-plane motions typically associated with sideways falls ͓23-25͔, we restricted our efforts in this preliminary study to two-dimensional models of backward falls. For this particular class of falls, our research questions were: ͑1͒ What attenuations in impact velocity and kinetic energy can theoretically be attained by the development of lower extremity joint torques during descent? ͑2͒ What is the theoretical effect on impact velocity and kinetic energy of declines in available joint torques? Methods The energy absorbed by a given joint during the descent phase of falling depends on the magnitude of joint torque and the magnitude of joint rotation. To determine how each of these variables theoretically affects impact severity, we developed one-link, twolink, and three-link inverted pendulum models of backward falls from standing height ͑Fig. 1͒. Among the assumptions inherent in these models were: ͑1͒ that movement is restricted to the sagittal plane, ͑2͒ that the feet remain stationary and in contact with the ground throughout descent, and ͑3͒ that contraction of muscles spanning the ankles, knees, or hips generates a net joint torque, which can instantly change in magnitude and direction. The onelink model simulates a fall where the knees and hips remain extended throughout descent, while rotation and energy absorption occur at the ankles. The two-link model simulates a fall where the knees are maintained in extension, while rotation and energy absorption occur at the ankles and hips. The three-link model simulates a fall where rotation and energy absorption occur at the ankles, knees, and hips. In each model, the lengths, masses, and moments of inertia of the various links were representative of an adult female of height 1.6 m and body mass 53.7 kg ͓26͔. Each model incorporated ideal torque generators to simulate the net effect of bilateral ͑equal right and left side͒ contraction of muscles spanning the ankles, knees, and hips. The effect of joint strength ͑or degree of muscle activation͒ on impact severity was determined by conducting simulations with different ''strength factors'' applied to each joint, ranging from zero to 100 percent of peak attainable values measured in young healthy females under isometric conditions ͑Table 1͒. For a given joint, the same strength factor was applied to flexor and extensor torques. However, in the two-link and three-link models, the strength factor of one joint was varied independently of that applied to the other joint͑s͒. For example, a three-link model simulation with strength factors of 75 percent at the ankles, 25 percent at the knees, and 50 percent at the hips would involve 150 Nm of ankle plantarflexor torque, 68 Nm of ankle dorsiflexor torque, 88 Nm of knee extensor torque, 39 Nm of knee flexor torque, 125 Nm of hip extensor torque, and 65 Nm of hip flexor torque. Based on experimental evidence of joint torque-rotation behavior during sits and self-initiated backward falls ͓22͔, the direction of joint torque always acted to raise ͑or retard downward motion of͒ the center of gravity of the link directly above the joint ͑Figs. 1 and 2͒. Consequently, in the two-link and three-link models, a reversal in the inclination of the trunk from forward to behind the vertical caused hip torque to change instantly from extensor to flexor, and vice versa. Similarly, in the three-link model, a reversal in the inclination of the shin from forward to behind the vertical caused ankle torque to switch abruptly from plantarflexor to dorsiflexor, and vice versa. It should be noted that, in real life, changes in the magnitude and/or direction of joint torques are gradual rather than instantaneous, due to finite rates of motor neuron recruitment and de-recruitment, and finite rates of change in muscle tension following neural activation. Each model was ''released'' from a configuration where the whole-body center-of-gravity was posterior to the ankle joint, and the total gravitational potential energy of the body ͑with respect to the ankles͒ equaled 410 J. Accordingly, the one-link model descended from an initial configuration involving 27 deg of plantar- In all simulations, initial angular velocities were set to zero. These initial conditions were selected to roughly describe those involved in a series of tether-release ͑falling͒ experiments conducted recently in our laboratory ͑currently unpublished͒. They are not necessarily intended to reflect typical body configurations after real-life slips or trips. MATLAB was used to numerically integrate the equations of motion ͑see Appendix A͒. At each integration step, the analysis routine computed translational and angular velocities of each link, and corresponding magnitudes of whole-body kinetic energy ͑see Appendix B͒. Modifications to MATLAB's ordinary differential equation solver were made to accommodate reversals in the direction of joint torque during a given simulation. Each simulation proceeded until the occurrence of impact, signified by the vertical position of the pelvis descending below the ankles. The change in potential energy (⌬ PE) during descent was defined as ⌬ PEϭ PE i Ϫ PE f , where i and f represent initial and final states. Our previous studies indicate that during an unexpected backward fall, individuals maintain the knees and hips moderately flexed during descent, and impact the ground with the trunk in a near-upright configuration ͓23͔. Therefore, to focus our attention on realistic falling scenarios and alleviate the need to incorporate joint ''stops'' into the model, we disregarded simulations involving initial impact to the knees or head, or the occurrence of hyperflexion or hyperextension at the knees or hips. The work performed at a given joint was determined by numerically integrating the area under the torque-rotation curve. Joint work was defined positive if the direction of torque was opposite to the direction of joint rotation ͑i.e., eccentric͒. Checks were made to ensure that conservation of energy was maintained throughout all simulations, as defined by ⌬ PEϭKE tot ϩW tot , where W tot is the sum of joint work, and KE tot is the total ͑translationalϩrotational͒ kinetic energy of the body at the instant of pelvis impact. Impact severity was represented by the vertical component of the body's total kinetic energy (KE v ) and the vertical ͑downward͒ , and 100 percent ͑no impairment͒. For simplicity, ''best-case'' falls were those that minimized KE v . Attention was also focused on impact severity during simulations with the two-link and threelink models involving zero ankle torque, which simulated falls from a narrow base of support. Results Impact severity was reduced substantially by the generation of lower extremity joint torques during descent. The best-case fall with the one link model, which involved a strength factor of 100 percent at the ankles, resulted in attenuations of 24 percent in KE v and 13 percent in v v ͑Table 2͒. The best-case fall with the twolink model, which involved strength factors of 100 percent at the ankles and 23 percent at the hips, resulted in attenuations of 62 percent in KE v and 21 percent in v v ͑Table 3͒. The best-case fall with the three-link model, which involved strength factors of 81 percent at the ankles, 23 percent at the knees, and 81 percent at the hips, resulted in attenuations of 79 percent in KE v and 48 percent in v v ͑Table 4͒. Comparison of outcome parameters from the three models indicates that rotation ͑and torque generation͒ at each joint influenced fall severity ͑Fig. 4͒. For example, comparison of best-case falls with the one and two link models suggests that rotation and torque generation at the hips reduces KE v by up to 50 percent ͑from 267 to 133 J͒ and v v by up to 10 percent ͑from 2.80 to 2.53 m/s͒. This is due to two phenomena. First, by generating eccentric extensor torque at the hips during descent, W tot increased from 99 to 141 J. Second, by flexing the hips to impact the ground with the trunk in a near-upright configuration, ⌬ PE decreased from 410 to 306 J. Comparison of best-case falls with the two and three-link models suggests that rotation and torque generation at the knees can attenuate KE v by a further 44 percent ͑from 133 to 74 J͒, and v v by a further 34 percent ͑from 2.53 to 1.68 m/s͒. Rather than being due to differences in W tot , ⌬ PE, or KE tot , this arose mainly from knee extension during the final stage of descent, which allowed for a ''transferring'' of energy from the vertical to horizontal direction, and a corresponding increase from 19 to 91 J in the horizontal kinetic energy of the body at impact (KE h ). Finally, comparing best-case simulations with the three-link model involving zero and 100 percent available strength factors at the ankles indicates that ankle torque can increase W tot and KE h by up to 34 and 60 percent, respectively, and decrease KE v and v v However, when compared to our worst-case fall, substantial reductions in impact severity were observed even with large declines in joint strength. For example, with strength factors of 19 percent at the ankles, 10 percent at the knees, and 19 percent at the hips, the three-link model yielded KE v ϭ178 J and v v ϭ2.81 m/s, attenuations of 50 and 13 percent, respectively ͑Table 4͒. Several mechanisms were responsible for this. First, even when accompanied by relatively small magnitudes of joint torque, large joint rotations allowed for nonnegligible magnitudes of W tot . Second, simulations with the two-link and three-link models continued to involve impact with the trunk in a nearly upright configuration, and a corresponding large reduction in ⌬ PE. Finally, optimal strength factors at the hip in the two link model and knee in the three link model were always no greater than 30 percent, and were therefore affected little by simulated declines in strength. Discussion Our results indicate that substantial reductions in fall severity can be achieved through the development of lower extremity joint torques during descent. For example, when compared to our worst-case fall, best-case falls with the three-link model resulted in attenuations of 79 percent in KE v and 48 percent in v v . Among the mechanisms responsible for this were: ͑1͒ the generation of eccentric ͑braking͒ torques in the lower extremity joints during descent, which resulted in up to 150 J of energy absorption; ͑2͒ the occurrence of impact to the ground with the trunk in an upright configuration, which reduced the change in potential energy during descent by 100 J; and ͑3͒ the occurrence of knee extension during the final stage of descent, which transferred up to 90 J of impact energy from the vertical to horizontal direction. Our results also suggest that one's ability to land safely depends at least as much on ''technique'' as it does on strength. For example, reductions of up to 50 percent in KE h were observed even with simulated declines of 80 percent in peak joint torques. Thus, properly executed ''relaxed'' falls should involve considerably lower risk for injury than rigid falls involving minimal knee and hip flexion ͑which might arise from fear or limitations on joint flexibility͒. Further evidence for the importance of technique in safe landing is the observation that best-case falls with the two-link and three-link models did not involve maximum torque activation. For example, best-case falls with the two link model involved relatively small magnitudes of hip torque ͑strength factors of approximately 20 percent͒; larger values caused the trunk to impact in an inclined position, which increased ⌬ PE and decreased W tot . Bestcase falls with the three-link model involved large hip torques but small knee torques ͑strength factors of approximately 20 percent͒, which facilitated large knee and hip flexions and thus large W tot . Competing biomechanical phenomena influenced the effect of joint rotation on impact severity. Knee and hip flexion tended to reduce KE v by increasing W tot and decreasing ⌬ PE. However, their effect on v v was more complex. On the one hand, such flexions caused the body's center-of-gravity to move closer to the ankles. This reduced the body's ''effective'' moment of inertia and, due to conservation of angular momentum, increased its effective rotational velocity ͑with respect to the ankles͒. This accounts for the relatively large magnitudes of v v predicted by the two-link model. On the other hand, knee flexion reduced the distance between the pelvis and ankles, and therefore the magnitude of v v for a given rotational velocity. This explains why v v was relatively small in best-case falls with the three-link model. Our simulations complement experimental evidence regarding the effect on fall severity of energy absorption in the lower extremity during descent. In a recent experimental study of backward falls with young subjects, we found that hip impact velocities averaged 1.45Ϯ0.5 ͑S.D.͒ m/s and vertical kinetic energies at impact averaged 31.6Ϯ24 J ͓22͔. These values are lower than our current predictions, probably due to the fact that our models descended from an initial configuration of imbalance, while subjects in the experimental study self-initiated their descent from a stable standing position, and absorbed a substantial amount of energy in their lower extremity joints before reaching a state of imbalance. In an earlier study ͓23͔, we found that, in addition to absorbing energy during descent, the responses elicited during unexpected falls serve to arrange the body in a safe landing configuration. For example, subjects tended to avoid impact to the head and pelvis by impacting the ground with the outstretched hands ͑and in the case of forward or sideways falls, the knees͒, and by rotating of the trunk about an inferior-superior axis ͑during sideways falls͒. However, backward falls in that study involved qualitatively similar joint rotations to those observed here, and while pelvis impact velocities were higher ͑averaging 2.55Ϯ0.85 m/s͒, they were again well below values predicted by free fall assumptions. Data also suggest that impact severity during sideways falls may be reduced by the generation of flexion rotations and eccentric extensor torques at the knees and hips during descent. The strongest evidence of this comes from van den Kroonenberg and co-workers' ͓23-25͔ study of body movements during selfinitiated sideways falls. Their reported hip impact velocities averaged 2.75Ϯ0.42 ͑S.D.͒ m/s, and their estimates of kinetic energy at impact averaged 188 J, or 71 percent lower than subjects' potential energy during standing. They also found that, when subjects were instructed to ''fall as relaxed as they could'' as opposed to ''naturally,'' reductions occurred in average values of trunk angle at impact ͑14 deg versus 22 deg from the vertical͒ and hip impact velocity ͑2.66 m/s versus 2.86 m/s͒. One factor apparently contributing to the latter observation was greater knee flexion during descent, and a subsequent reduction in the distance between the pelvis and ankles ͑which, as observed in our three link model simulations, reduces the translational velocity of the pelvis for a given rotational velocity͒. Several limitations exist to this study. We examined only a single torque activation strategy ͑i.e., attempting to raise the link above the joint͒, and while preliminary experimental results suggest this to be realistic ͓22͔, alternative and potentially more effective muscle activation strategies may exist. We also neglected cases involving hyperextension or hyperflexion at the knees or hips, which might be examined through realistic simulation of joint stops. Moreover, we simulated the net effect of muscle contractions about a joint with ideal torque actuators, and did not account for the effect on torque development of variables such as the intactness of proprioceptive and vestibular signals, the state of potentiation and firing frequency of motor neurons, the intrinsic ͑force-length and force-velocity͒ properties of muscle and inseries connective tissue, the anatomical arrangement of muscles spanning each joint, and the degree of co-contraction of agonist and antagonist muscles. We doubt that adding these features to the model would substantially change our conclusions regarding the effect on impact severity of available torque magnitudes. However, it would provide a more robust tool for systematically examining how fall severity is influenced by specific neuromuscular pathologies. We also assumed that the feet remain fixed on the ground and that dorsiflexor ͑or plantarflexor͒ moments can be generated at the ankles throughout descent. During an actual fall, the toes or even both feet may rise off the ground. Falls involving complete loss of foot-ground contact ͑due, for example, to a violent slip or trip͒ would likely be severe, since it is ultimately the development of vertical foot reaction forces that decelerates the body's downward movement. The fact that we did not restrict plantarflexion rotation probably caused us to overestimate energy absorption at the ankle in simulations with the one link and two link models, where the ankle rotated to an unrealistic 90 deg of plantarflexion before

Cardiovascular Responses to Orthostasis and Their Association With Falls in Older Adults

Background Orthostatic hypotension (OH) refers to a marked decline in blood pressure when upright. OH has a high incidence and prevalence in older adults and represents a potential intrinsic risk factor for falls in these individuals. Previous studies have not included more recent definitions for blood pressure responses to orthostasis, including initial, delayed, and recovery blood pressure responses. Furthermore, there is little research examining the relationships between cerebrovascular functioning and falling risk. Therefore, we aimed to: (i) test the association between different blood pressure responses to orthostatic stress and retrospective falling history and; (ii) test the association between cerebrovascular responses to orthostatic stress and falling history. Methods We tested 59 elderly residents in long term care facilities who underwent a passive seated orthostatic stress test. Beat-to-beat blood pressure and cerebral blood flow velocity (CBFV) responses were assessed throughout testing. Risk factors for falls and falling history were collected from facility records. Cardiovascular responses to orthostasis were compared between retrospective fallers (≥1 fall in the previous year) and non-fallers. Results Retrospective fallers had larger delayed declines in systolic arterial pressure (SAP) compared to non-fallers (p  = 0.015). Fallers also showed poorer early (2 min) and late (15 min) recovery of SAP. Fallers had a greater decline in systolic CBFV. Conclusions Older adults with a positive falling history have impaired orthostatic control of blood pressure and CBFV. With better identification and understanding of orthostatic blood pressure impairments earlier intervention and management can be implemented, potentially reducing the associated risk of morbidity and mortality. Future studies should utilize the updated OH definitions using beat-to-beat technology, rather than conventional methods that may offer less accurate detection