Anatomical hip model for the mechanical testing of hip protectors

Medical Engineering & Physics 27 (2005) 475–485

Anatomical hip model for the mechanical testing of hip protectors Siegfried Derler a,∗ , A.B. Spierings a , K.-U. Schmitt b a

Swiss Federal Laboratories for Materials Testing and Research (EMPA), Lerchenfeldstrasse 5, CH-9014 St. Gallen, Switzerland b Institute for Biomedical Engineering, Swiss Federal Institute of Technology (ETH) and University of Zurich, Gloriastrasse 35, CH-8092 Z¨urich, Switzerland Received 9 September 2004

Abstract An anatomical hip model has been developed to simulate the impact load on the hip of a falling person wearing a hip protector. The hip consists of an artificial pelvis made of aluminium, linked by a ball-and-socket joint to an anatomically shaped steel femur (thigh bone). The femur is embedded in silicone material with a hip-shaped surface to allow realistic positioning of the protectors with accessory underwear. Additionally, the silicone simulates the damping and load-dispersal effect of soft tissue. A triaxial load sensor is integrated in the neck of the femur to measure the axial and cross-sectional force components in response to external impact forces on the hip. The performance of the hip model was investigated in drop tests and validated against biomechanical data. In a first series of measurements, the shock absorption of 10 different hip protectors, including both energy-absorbing and energy-shunting systems, was analysed. To determine the importance of hip protector placement, each protector was tested in the correct anatomical alignment over the hip and anteriorly displaced by 3 cm. Considerable differences were found between individual hip protectors in their effectiveness to reduce impact forces on the femur. Position of the hip protector also influenced the forces applied to the femur. © 2005 IPEM. Published by Elsevier Ltd. All rights reserved. Keywords: Hip protector; Anatomical hip; Femur model; Impact load; Hip fracture

1. Introduction Hip fractures represent a severe health problem for the elderly. In many countries, large increases in hip fracture incidence are expected due to increasing life expectancy and ageing populations [1–6]. In Switzerland, the current annual incidence rate of hip fractures among elderly people is slightly above 0.1%, resulting in approximately 8000 hip fractures per year [7]. In most cases, hip fractures are the consequence of falls, which typically occur from standing height towards the side, and are characterised by a direct impact to the hip [8–16]. According to Kannus et al. [17], 25% of sideways falls in the elderly cause hip fractures. One possibility to prevent hip fractures is the use of hip protectors. At present, mainly two types of products, namely energy-absorbing soft pads and energy-shunting hard shells, ∗

Tel.: +41 71 274 77 66; fax: +41 71 274 77 62. E-mail address: [email protected] (S. Derler).

as well as combined systems, are used to protect the hip in the case of impact loads in the area of the greater trochanter. The principle of an energy-shunting hip protector is to distribute impact loads away from the trochanter to the surrounding soft tissue, while an energy-absorbing device attenuates impact forces by means of a shock-absorbing material. In 1988, Wortberg proposed the first hip protector, which was made of a special silicone rubber [18]. Since then, alternative concepts such as the use of airbags [19] or a fluid-containing pad system [20] have also been suggested. The effectiveness of hip protectors has been investigated previously by means of mechanical testing and clinical case–control studies. Even though the majority of clinical trials reported positive results on the use of hip protectors [2,21–26], critical remarks concerning not only acceptance and proper usage [27,28], but also the effectiveness of hip protectors can be found in some recent studies [29,30]. Because each of the clinical studies focused on one specific device and differed in detail, direct comparisons between the results and general conclusions are problematic [31].

1350-4533/$ – see front matter © 2005 IPEM. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.medengphy.2005.02.001

476

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

Mechanical testing represents an important, complementary method to assess the efficacy of hip protectors both objectively and economically, and to support the development and optimisation of protector materials and designs. In a number of previous studies, the impact conditions of the hip in a sideways fall were experimentally simulated based on biomechanical data [20,32–34]. Previous testing systems have used different measurement principles (either drop tests or impact pendulum tests) and test parameters have varied widely in the degree to which the mechanical properties and the geometry of the human hip were modelled. In the impact pendulum tests of Parkkari et al. [32] and Robinovitch et al. [20], the stiffness and the damping of the pelvis were simulated by means of springs and bumpers, respectively, while polyethylene foam was used to mimic trochanteric soft tissue. In contrast, elastomers served as soft tissue surrogates on a rigid base in the drop tests of Mills [33] and Nabhani and Bamford [34]. Robinovitch et al. [20] produced a hip form with a realistic surface geometry, whereas Mills [33] used a cylindrical shape, and Nabhani and Bamford [34] constructed only a small hip section. The previous methods had in common that forces, transmitted to an artificial greater trochanter, were measured using a uni-axial load cell in order to assess the force attenuation capacity of hip protectors in impact tests. The main disadvantage of a uni-axial load cell is that only one projection of the transmitted force is measured, while oblique force components, e.g. those transmitted to the femoral neck over the femur shaft, cannot be resolved. None of the testing systems used so far allowed the realistic simulation of practical wearing conditions for hip protectors. In addition, geometrical simplifications of the femur or surrounding soft tissue might influence the distribution of impact forces. Thus, mechanical testing methods need further development. According to Cameron and Kurrle [35], appropriate testing and regulation is not only required for hip protectors themselves, but also for underwear and other systems used to wear the protector. The objective of the present study was to develop an anatomically realistic mechanical hip model, on which hip protectors can be positioned and impact-tested under biomechanical conditions, which are typical for falls in the elderly.

Besides the anatomically shaped structures of the hip model, the tri-axial measurement of forces in the femoral neck for the assessment of hip protectors represents a major improvement compared to previous methods. In the following, details on the design of the hip model are presented and the function of the apparatus is characterised on the basis of first measurement series using the hip model alone as well as with hip protectors.

2. Experimental design An anatomically realistic mechanical hip model was developed, on which all types of hip protectors and accessory underwear can be placed correctly or out of position for impact testing. To minimize geometrical restrictions, the left side of a human torso ranging from the knee to the waist area was constructed. A flesh surrogate with a realistic thickness above the greater trochanter extends over a region wider than is covered by current hip protectors, in order to provide a realistic distribution of applied impact forces. An anatomically shaped femur model is embedded in the flesh surrogate. It incorporates a tri-axial load cell in the femoral neck, allowing axial and cross-sectional force components in the area of cervical hip fractures to be measured and analysed for attenuation and distribution of impact forces. Because it is known that hip fractures are twice to three times as frequent for women as for men [36], the female anatomy was used as an overall design criterion for the hip model. 2.1. Design and geometry of the hip model Fig. 1 shows a schematic of the mechanical hip model for simulating the impact loads on the hip in a sideways fall. The geometry of the femur was derived from a standard femur model, which was proposed by Viceconti et al. [37] for biomechanical finite element analyses. The 3D surface of this standard femur was approximated using a CAD program, thereby taking special care to capture the important anatomical features of the greater trochanter. Geometrical simplification of the femoral neck region was necessary to

Fig. 1. Frontal view of the anatomical hip model (axis labels in cm).

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

Fig. 2. Artificial hip joint and proximal femur.

integrate a load cell. Fig. 2 shows a close-up of the proximal part of the resulting femur model (steel), which is linked to a simplified pelvis via a ball-and-socket joint. The femur has a length of approximately 44 cm, which lies in the range of the results of Mall et al. [38], who investigated the geometry of the femora in a German population of 100 men (46.4 ± 2.4 cm, mean age 61 years, mean height 171 cm) and 70 women (43.4 ± 2.4 cm, mean age 72 years, mean height 161 cm). The model is further characterised by a femur head diameter of 4.5 cm, neck and shaft diameters of 2.7 cm, a femur axis length of 10.2 cm and a neck–shaft angle of 127◦ . These parameters are comparable to the average dimensions of female femora reported by Mall et al. [38]. The artificial pelvis was modelled on the female anatomy. Geometrical measurements were carried out on a skeleton model of medium size (Somso, female pelvis with femora) in order to design a simplified pelvis with realistic hip joint and iliac crest, over which the upper part of a hip protector could be extended, depending on the specific design and positioning. With an inclination of 10◦ to the horizontal and a posterior rotation of 12◦ , the orientation of the femur was defined close to the parameters used by Courtney et al. [39] and Pinilla et al. [40] to investigate the fracture loads of isolated femora. This geometrical configuration is associated with the impact of the hip in a sideways fall [40] and fracture patterns of isolated bones were found to be comparable to real cervical hip fractures [41]. The distal end of the femur is connected to an axial knee joint in order to prevent rotation of the leg about the hip joint. The knee joint can take up forces perpendicular to the femur shaft, torsional moments, as well as moments in the saggital plane. The rotational motion of a human leg about the hip joint due to impact forces on the hip was estimated by applying Newtonian mechanics to a rigid beam, which freely rotates about a fixed point. Assuming a moment of inertia of 1.8 N m2 for this leg model (based on leg dimensions and masses given in [42]) and an impact force comparable to the experimental results for the hip model (see Section 3.1), negligible rotation of the leg (<2◦ ) was found throughout the entire impact force pulse. Little is known about friction forces in the human hip joint, but frictional coefficients are assumed to be very low in

477

healthy hip joints. Values of 0.05 are typical for artificial hip joints [43]. Since muscles are likely to be activated in a fall to the side [44] additional forces and torques can be generated that hinder the rotation of the leg. To take into account the effects of friction and muscle forces in the mechanical hip model, a moment which is necessary to rotate the artificial leg within the coronal plane, has been specified for the hip joint (an arbitrary value of 17.7 N m was defined as no data was found in the literature). This (adjustable) moment is given by friction in the ball-and-socket joint in combination with the mass and stiffness of the materials used for bone and soft tissue. In a computational study, in which the moment was varied between 0 and 35 N m, we found that the force response of the hip model was changed by less than 2.5%. To define the anatomical shape of the flesh surrogate, surface data of the hip region of a human figure [45] were fitted to the femur/pelvis combination, thereby using simple scaling functions and taking into account the anthropometric 50thpercentile measures given by J¨urgens [46]. 2.2. Materials The femur model was made of steel and the artificial pelvis was made of aluminium. Steel was used instead of a bonelike material because of its robustness. For both steel and real bones, large deformations, i.e., substantial geometrical changes and influences on the force transmission, can be excluded for forces below the fracture load. The same applies to aluminium, which was chosen to minimize the weight of the apparatus. It is important to note that the metal structures of the mechanical hip simulate the force transmission, but not the deformation behaviour of human hip bones. For this purpose, the selected materials have to fulfill requirements concerning rigidity, without necessarily matching the material properties of bones. To reduce friction between the femur head and the pelvis, the surface of the hip joint socket was coated with an aluminium–polytetrafluoroethylene (PTFE) compound. The pressure between ball and socket of the joint can be varied to adjust the moment acting against the rotation of the femur about the hip joint. The selection of the flesh surrogate is of crucial importance, since it is known that the layer of soft tissue in the region of the greater trochanter can absorb a significant portion of the impact energy and thus strongly reduce the impact forces [47]. In previous test methods, materials such as rubber sheets with a thickness of 10–50 mm [48], 20–25 mm-thick polyethylene foams [17,20,32], a 20 mm-thick polymer with a density of 1110 kg/m3 [33] and a 5 mm-thick silicone elastomer [34] were used to simulate the influence of soft tissue. To determine the material properties of soft tissue in the area of the greater trochanter, compression tests were carried out on the hips of 10 human subjects using an Instron 4502 apparatus. Force–deformation curves were measured up to a load of 200 N at a deformation rate of 10 mm/min and compared to the results of experiments on silicone elastomer samples with varying composition. A silicone elastomer with

478

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

a density of 1230 kg/m3 (Wacker Elastosil M 4511), which showed a material behaviour similar to relatively stiff human hips at quasi-static conditions, was chosen as flesh surrogate. A thickness of 20 mm was defined for the layer above the most prominent part of the greater trochanter, corresponding to a typical thickness of soft tissue found for female hip-fracture patients [49]. The silicone elastomer was moulded around the metallic components of the proximal femur. It extends from the thigh to the waist, reaching a maximum thickness of 70 mm. The rest of the hip form was made of ethylene propylene diene monomer (EPDM)-foam, allowing the realistic simulation of wearing conditions for hip protectors and underwear. To protect the materials of the hip model against surface abrasions, they were covered with a compression hose fabric. 2.3. Impact tests and instrumentation The hip model with a positioned protector is used in a drop test as shown in Fig. 3. For the impact loading of a hip protector, a mass with a flat impact surface (radius 10 cm) falls from a pre-defined height. The impact point can easily be varied, but normally the center of the falling mass is vertically aligned above the greater trochanter. Hip protectors can be placed either correctly on the hip or out of position to simulate displacements occuring in practice. In the stan-

dard experiment, a falling mass of 5, 10 or 15 kg is dropped from a height of 50 cm, measured from the surface of the hip model without protector. This results in an impact velocity of 3.1 m/s and in impact energies of 24.5, 49.1 or 74.6 J, respectively. The specified velocity is a typical value for the impact velocity of the hip on the floor resulting in a sideways fall [50–52]. A uni-axial accelerometer (Kistler Type 8005) was used to measure the deceleration of the falling mass during impact. From the deceleration signal, the impact force acting both on the falling mass and on the hip protector is calculated. In addition, the measured deceleration is integrated to determine the deformation of the hip/protector system. The force transmitted to the femoral neck was measured by a tri-axial force transducer (Kistler Type 9047B), providing both axial and cross-sectional force components for further analysis. The load cell was placed in the femoral neck, because this represents a critical area in which cervical hip fractures often take place [4,53]. Its position, which is indicated in Fig. 2, corresponds to the femoral neck region where the highest stresses are found [54]. Due to the incorporated load cell, the geometry of the femoral neck deviates from the real anatomy (see Fig. 2). These geometrical deviations are expected to have a negligible effect on the forces measured in the femoral neck, as external impact loads are mainly transmitted via the greater trochanter and the femur shaft. The signals of the accelerometer and the force transducer were filtered using a 1 kHz low-pass filter (CFC 600 according to ISO 6487: 1987)1 and sampled at a rate of 100 kHz over a period of 40 ms using a PC equipped with a transient recorder board. High-speed video analysis was used to study the deformation behaviour of the bare hip model in impact tests.

3. Model performance In this section, the impact response of the mechanical hip model is characterised in detail. This not only allows a comparison with previous test methods and the results of biomechanical studies on falls and hip fractures, but also serves as a basis for the interpretation of the hip protector test results presented below. 3.1. Impact response of the hip model Fig. 4 shows a typical impact response of the bare hip model in a drop test using a mass of 10 kg falling from a height of 50 cm (impact energy 49.1 J). The impact force is characterised by a peak value of 7.1 kN and a time-to-peak of 8.3 ms. Despite a slightly different pulse shape, the resultant neck force starts almost simultaneously with the impact force and shows a comparable rise time of 8.1 ms to

Fig. 3. Drop test for hip protectors.

1 ISO 6487: 1987 Road vehicles – Measurement techniques in impact tests – Instrumentation, International Organisation for Standardization

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

Fig. 4. External impact force and femoral neck forces resulting in a drop test on the hip model alone (using a mass of 10 kg falling from a height of 50 cm).

the peak value of 5.7 kN, which is about 20% lower than the maximum impact force due to the damping effect of the flesh surrogate as well as for geometrical reasons discussed below. The maximum impact force measured with the anatomical hip model was within the range of biomechanical results on hip impacts [51], and exceeded the threshold value for femoral fracture, which is reported to be 2.5 kN [34]. The hip impact forces in a sideways fall depend on a variety of factors such as muscle tension, energy absorption occurring by soft tissue, protective responses of the falling person, effective (moving) mass of the hip, impact configuration, impact velocity and floor covering. Robinovitch et al. [47] placed isolated tissue samples with varying thickness on an artificial femur for impact testing (140 J) and measured impact forces between 4.05 and 6.42 kN as the thickness of the soft tissue decreased from 43 mm to 8 mm. With the mechanical hip model, higher impact forces are measured at lower impact energy, i.e., the flesh surrogate, which has a thickness of 20 mm above the greater trochanter, is considerably stiffer than human tissue. This is also evident from the relatively short duration of the impact force pulse shown in Fig. 4. In biomechanical experiments, in which a person fell to the side from standing height, Askegaard and Lauritzen [50] found rise times of around 30 ms in the hip impact forces. For subjects who carried out sideways falls from a kneeling position, Sabick et al. [55] measured typical rise times of about 70 ms in the hip impact forces. In both studies, however, the impact force pulses were attenuated and broadened by the use of shock absorbing foam materials (in addition to the effects of biological tissues). Parkkari et al. [56] impacted the hips of young volunteers who wore hip protectors on both sides of the pelvis and stood against a concrete wall. For impact energies around 90 J, typical impact force pulses showed rise times around 25 ms. For the unprotected hip, the rise time of the impact force pulse might have been considerably shorter due to the fact that protectors generally increase the available distance to decelerate an impacting mass. Since impact tests on unprotected human hips are not feasible at energies bearing an injury risk, however, it is not known how

479

short impact force pulses can be in reality. Owing to the firm flesh surrogate combined with rigid metal bones, we consider the mechanical hip model to simulate a stiff human hip and, thus, to provide critical test conditions for hip protectors. In a standard drop test on the hip model alone, the flesh surrogate absorbs an energy of 29.1 ± 0.7 J, as was determined from force–deflection curves. This lies in the range of tissue energy absorption (34 ± 26 J) reported by Robinovitch et al. [47] and corresponds to 59% of the applied impact energy of 49.1 J. In contrast to the human hip, the mechanical model can absorb energy only in the flesh surrogate, but not in the steel femur or in the aluminium pelvis. In quasi-static experiments, Beason et al. [57] found a mean fracture energy of 75 ± 33 J for the pelvis, whereas for proximal femora of the elderly mean fracture energies between 22 and 83 J have been reported [39,58,59]. By combining these different results, the energy absorption of pelvis, femur and soft tissue together can be estimated between about 70 and 250 J. Since the fracture energy of the pelvis is normally higher than that of the femur, however, the pelvis is expected to absorb only a part of its fracture energy in the case of a femur fracture. Considering the differences between a real hip and the mechanical hip (no energy absorption of femur and pelvis), it can be assumed that an experiment in which an impact energy of 49.1 J is used is comparable to cases in which the human hip is loaded with impact energies higher than 100 J. The total energy absorbed by the human hip is much lower than the potential energy of a falling person, which is typically around 500 J [33,52,60,61]. The range of experimental impact energies chosen in this study (up to 74.6 J) is comparable to previous investigations on hip protectors [17,32,33,62]. Robinovitch et al. [20] and Nabhani and Bamford [34] used an impact energy of 120 J, while Kannus et al. [17] and Parkkari et al. [32] also carried out impact tests at 110 and 132 J, respectively. Except the drop tests of Nabhani and Bamford [34], impact energies above 100 J were only used in pendulum tests and in connection with an experimental set-up that simulated the stiffness and the energy absorption of the human pelvis. Measured impact forces and forces transmitted to the femur of the anatomical hip model can be compared to the fracture loads of human femora. In quasi-static experiments (loading rate 100 mm/s), Pinilla et al. [40] and Courtney et al. [39] investigated isolated femora in a geometrical configuration (femur shaft inclined 10◦ to the horizontal, internal rotation of 15◦ ), which is comparable to that of drop tests using the anatomical hip model. Pinilla et al. [40] found fracture forces of 3.82 ± 0.91 kN for 81 ± 7 years old femora, while Courtney et al. [39] reported fracture loads of 4.16 ± 1.59 kN for 73 ± 8 years old bones. The applied loads corresponded to vertical forces on the greater trochanter. In both studies, the mean values minus one standard deviation were close to the threshold value for fracture of the femur mentioned above (2.5 kN). Since the development of an alternative

480

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

criterion lies outside the scope of the present study, the same value is used here as a provisional limit, allowing a first interpretation of the impact test results obtained with the hip model. It has to be kept in mind, however, that forces exerted directly on the greater trochanter are different from both external forces on the hip and forces in the femoral neck. It follows from the geometry of the hip model that a vertical force on the greater trochanter of the artificial femur causes a resultant force in the femoral neck which is about 91% of the external force. 3.2. Measurement uncertainty and sensitivity The repeatability of the impact response of the hip model was investigated in measurement series of 10 repetitive impact tests. For the peak values of the impact force, the coefficients of variation were between 1% and 2%, whereas those of the resultant force in the femoral neck and of the individual force components ranged from 2% to 3%. Using the same moulded flesh surrogate, the loading response of the hip was reproduced for more than 200 impact tests. In the area of the greater trochanter, the surface of the hip has a radius of curvature of about 10 cm in the axial crosssectional plane and of about 50 cm in the longitudinal section, respectively. This curvature, in combination with the shape of the femur, makes the hip model sensitive to local variations in the impact point and changes in the distribution of impact loads. To quantify this sensitivity, a series of drop tests was carried out in which the impact point of the falling mass was varied in different directions from the normal impact point. Anterior and posterior displacements of 1 cm systematically reduced the measured impact forces between 1% and 2%, and the resultant femoral neck forces between 2% and 3%. While inferior and superior displacements of 1 cm had practically no effect on the impact force, the resultant forces in the femoral neck increased by 7% for inferior and decreased by 8% for superior displacements, respectively. As expected, the transmission of impact forces to the femoral neck is especially effective for geometrical configurations in which the force distribution involves the femur shaft area. In drop tests, the impact point is hit within deviations of 2–3 mm. The resulting contributions to the measurement uncertainty of forces are therefore less than 2% for displacements of the impact point in all directions.

4. Hip protector tests The mechanical hip was used to investigate 10 different hip protector models, including three energy-shunting systems with hard shells and seven energy-absorbing designs, consisting of soft foam pads. Selected results of the first test series are reported here to illustrate the function and the potential of the anatomical hip model, whereas a detailed analysis of the effectiveness of hip protectors will be presented in a separate paper.

Table 1 Maximum impact forces (Fimpact ) and maximum resultant femoral neck forces (Fneck ), measured for 10 different protectors using impact energies of 24.5 J (a), 49.1 J (b) and 74.6 J (c) Protector

Maximum impact force Fimpact (N) a

b

c

a

b

c

A B C D E F G H I J

2263 2462 2417 2535 2303 2275 2179 3007 2096 2381

4673 4333 4231 4795 4014 4150 3955 5625 3428 4215

7386 5971 5913 7358 5798 6926 5827 7803 5317 5877

1604 1659 1759 1641 1729 1472 1361 2279 1010 1817

3552 2591 2773 3131 2782 2696 2266 4182 2085 2900

5816 3402 3753 5150 3935 5194 3522 5706 3598 3883

Hip model

–

7097 ± 69

–

–

5654 ± 102

–

Maximum femoral neck force Fneck (N)

Whereas the results of two and three test series were averaged for the lowest (a) and the medium impact energy (b), respectively, the results for the highest impact energy (c) are based on a single test series. The last line of the table contains mean values and standard deviations of 20 impact tests (49.1 J) on the hip model alone.

4.1. Impact forces and resultant forces in the femoral neck Two samples of each protector model were tested using impact energies of 24.5, 49.1 and 74.6 J. With each sample, three impact tests were carried out on different days, so that the protector materials could recover between successive measurements. Results are given in Table 1. With an impact energy of 24.5 J, the measured maximum impact forces (Fimpact ) ranged from 2.1 to 3.0 kN. Doubling the impact energy led to values between 3.4 and 5.6 kN, and the three-fold energy to maximum impact forces between 5.3 and 7.8 kN. The resultant forces in the femoral neck (Fneck ) were systematically lower and ranged from 1.0 to 2.3 kN for the lowest, from 2.1 to 4.2 kN for the middle, and from 3.4 to 5.8 kN for the highest impact energy, respectively. As Table 1 illustrates, the middle impact energy of 49.1 kN was high enough to cause the majority of impact forces and femoral neck forces to lie above the threshold value 2.5 kN for fractures of the femur discussed in Section 3.1. The forces in the femoral neck showed a higher variation for the different protector types than the impact forces, i.e., were more sensitive to variations in the force distribution due to different hip protector designs and modes of function. For the impact energy 49.1 J, the results of hip protector tests could be compared to those of drop tests on the hip model alone (Table 1). With hip protectors, the maximum impact forces varied between 47% and 80% of the mean value found without protector. On the other hand, the resultant forces in the femoral neck were reduced to values between 31% and 67% of the mean value resulting without protector. From the fact that greater reductions were found in Fneck , it can be concluded that by means of protectors significant portions of the forces were deflected via the flesh surrogate. As can

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

481

Fig. 5. Maximum external impact forces (Fimpact ), and resultant femoral neck forces, Fneck , for 10 different hip protectors (mean values of three drop tests using a mass of 10 kg falling from a height of 50 cm).

be seen in Fig. 5, the two assessment criteria Fimpact and Fneck lead to different rankings of the hip protector models investigated. 4.2. Comparison between energy-shunting and energy-absorbing protectors In Fig. 6, the maximum resultant forces in the femoral neck are plotted as a function of the maximum impact forces (data from the measurement series discussed in Section 4.1). Least squares fitting of quadratic functions indicate that energy-shunting hip protectors tend to transmit higher loads to the femoral neck than energy-absorbing protectors at highimpact forces. A possible explanation is that a part of the impact force is focused on the femur shaft by the inferior rim of an energy-shunting protector. This effect might be less important at low impact energies because of a greater relative contribution of the flesh surrogate to the total shock absorption of the hip/protector system. Systematic differences between energy-shunting and energy-absorbing protectors also become evident from Fig. 7, in which rise times of Fimpact (Fig. 7a) and of Fneck (Fig. 7b) are plotted against maximum deformations of protectors in combination with the hip model. It is not directly possible

Fig. 6. Maximum resultant forces in the femoral neck (Fneck ) as a function of external impact forces (Fimpact ) for 10 different hip protectors (data of six measurement series using masses of 5, 10 and 15 kg falling from a height of 50 cm).

Fig. 7. Rise times of measured external impact forces, Fimpact (a) and resultant forces in the femoral neck, Fneck (b) as a function of maximum deformations (data of six measurement series using masses of 5, 10 and 15 kg falling from a height of 50 cm).

to separate the contribution of the hip model from that of a protector to the total deformation. For impact tests on the hip model alone, the maximum deformation of the flesh surrogate turned out to be about 15 mm (estimated from highspeed video analysis) for an impact energy of 49.1 J. Due to their vaulted shapes, energy-shunting protectors showed relatively large deformations (Fig. 7a and b). For comparable deformations, on the other hand, energy-absorbing protectors were characterised by systematically longer rise times of Fneck (Fig. 7b). 4.3. Importance of hip protector alignment In practical use, hip protectors can be placed incorrectly or become displaced out of the optimum position by the wearer’s movements. As a consequence, their protective function can be compromised. This problem was investigated by testing two new samples of each protector model using the impact energy 49.1 J. One sample was placed correctly on the hip model whereas the other was displaced by 3 cm in the anterior direction, thereby simulating a realistic out-of-position problem. Results of impact tests on displaced compared to normally positioned hip protectors are given in Table 2. The changes found in the resultant forces in the femoral neck Fneck were generally greater than those in the maximum impact forces Fimpact , indicating that Fneck is the more sensitive and therefore more reliable criterion to assess the performance of hip protectors. For seven hip protector models, the absolute

482

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

Table 2 Effect of hip protector displacements on the maximum impact force (Fimpact ) and on the maximum resultant force in the femoral neck (Fneck ) for 10 different protector models Protector

Effect on maximum Fimpact (%)

Effect on maximum Fneck (%)

A B C D E F G H I J

−1.7 2.1 0.1 −4.2 −2.6 0.7 −6.9 2.5 8.8 −2.6

1.9 4.7 3.7 −11.2 5.3 −4.2 −15.2 1.7 23.4 6.8

Of each type, one protector was placed correctly on the hip model, whereas the other sample was placed out of position (anterior displacement of 3 cm) prior to testing. All protectors were investigated using an impact energy of 49.1 J. The effect of the displacement on maximum forces is expressed as the percentage 100 × (Fdisplaced − Fnormal )/Fnormal .

changes caused by displacements were below 7% in Fneck and below 3% in Fimpact ; in three cases, the absolute changes in Fneck lay between 11% and 23% and in Fimpact between 4% and 9%. Interestingly, the measured forces were not generally higher when the hip protectors were displaced. Fig. 8 shows a comparison between two specific types of protectors, for which the results were analysed in detail. With anterior alignment of the energy-shunting protector the impact force increased 9%, accompanied by an increase in the resultant force in the femoral neck of 23% (black curves in Fig. 8a and b). Thus, the level of protection was significantly reduced by the displacement of the protector. Conversely, reduced im-

pact forces (−4%) and lower resultant forces in the femoral neck (−11%) were found when an energy-absorbing protector was tested in an anterior position (gray curves in Fig. 8a and b). This implies that the design of this protector and/or its placement in the underwear is suboptimal. The examples presented in Fig. 8 indicated that the protective function of hip protectors is greatly influenced by correct alignment and design. 4.4. Orientation of the resultant force in the femoral neck Since force components in the femoral neck are measured in three axes, the spatial orientation of the resultant force can be determined in order to study geometrical effects on the impact dynamics. The orientation of Fneck depends on the distribution of Fimpact , which is given by the dynamics of the falling mass during impact. Therefore, Fneck is influenced by the protector design and the effective curvature of the hip/hip protector combination. Fig. 8c shows the two angles θ and φ of the spherical coordinates of Fneck during impact for the two examples of hip protectors discussed in Section 4.3. θ is the angle measured from the positive z-axis, and the angle φ is measured in the x/y-plane from the positive x-axis (the axes x, y and z are shown in Fig. 2). With the correctly placed protectors that were impacted at the centre, the orientations of Fneck were qualitatively comparable for the energy-shunting and the energy-absorbing protector, respectively. Both results indicate a slight rotation of the falling mass around an anterior–posterior axis, corresponding to a rolling in the longitudinal section of the hip model. The differences in the angle θ can be attributed to different effective radii of curvature

Fig. 8. Forces Fimpact (a), Fneck (b) and orientation of Fneck (angles θ and φ of spherical coordinates, see text) during impact (c) found for two different hip protectors in the correct position and anteriorly displaced. The points in plot (c) indicate the orientation of Fneck at the moment of peak force. At this moment, the orientation of Fneck roughly corresponds to the vertical direction, but in detail depends on the individual hip protector and the alignment.

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

of the two protectors in combination with the hip during impact. When the energy-absorbing protector was shifted from the normal position, the total rotation of the impacting mass was reduced, and the impact forces were attenuated, possibly due to a better distribution over a larger area. In contrast, the displacement of the energy-shunting protector led to a considerable rotation of the impacting mass around a superior–inferior axis. This means that the falling mass compressed a peripheral region of the hip protector and probably contacted the partly unprotected hip model. As a consequence, the peak forces of Fimpact and Fneck increased significantly. Further analyses will clarify the relationship between the performed drop tests and impact configurations occurring in practice. However, our first results on an energy-shunting hip protector indicate that the effectiveness of this specific design could be problematic if the protector is misplaced or pushed aside during an impact.

5. Summary and outlook Compared to previous testing systems, the presented mechanical hip model is characterised by a number of new features, allowing for improved testing and analysis of the efficacy of hip protectors: • The torso with anatomical hip geometry allows the realistic placement and impact testing of all types of hip protectors with the accessory underwear both in the correct position and out of the optimum position. • Due to its material properties and large dimensions, the silicone flesh surrogate moulded around the proximal femur provides a realistic deformation and force distribution of hip protectors under impact loads. • The artificial femur incorporates a tri-axial force sensor in the femoral neck to measure axial and cross-sectional force components in the area of cervical hip fractures. The analysis of the measured forces provides detailed information about the attenuation and distribution of impact forces by hip protectors. No previous testing system used a complete torso with anatomically shaped hip and femur for the investigation of hip protectors. In its current version, the hip model is used to simulate the impact loads on the hip resulting in a sideways fall. The orientation of the femur has been chosen to enable the forces measured in the femoral neck to be compared to the fracture loads determined for isolated femora [39,40], but the construction of the artificial hip joint allows the orientation of the femur to be varied. The orientation of the whole hip model can also be changed to simulate other loading configurations. The anatomical hip model showed good measurement repeatability and a high sensitivity to variations in the distribution of impact forces. Initial experiments using the new hip

483

model addressed the shock-absorbing behaviour of 10 different hip protectors as well as the problem of hip protectors, which were displaced from the optimum position. While clear differences were found between the individual hip protector models, no general advantage became evident for energyabsorbing or energy-shunting systems, respectively. Compared to seven models of energy-absorbing hip protectors, three investigated energy-shunting designs were characterised by greater maximum deformations and shorter rise times of the resultant femoral neck force. In addition, energyshunting hip protectors led to slightly higher loads of the femoral neck for high-impact forces. The inferior rim of an energy-shunting protector can cause a force concentration in the area of the femur shaft, resulting in an increased load on the femoral neck. For most of the investigated hip protectors, displacements out of the optimum position led to measurable changes in the forces transmitted to the femoral neck. Surprisingly, the forces were in some cases reduced. The fact that certain hip protectors seemed more effective when out of position indicated that either their design or their placement in the accessory underwear is suboptimal. The analysis of one specific energy-shunting hip protector compared to an energy-absorbing protector demonstrated that the orientation of the resultant force in the femoral neck can be used to study geometrical details of the impact of a falling mass onto a hip protector. In future work, this potential of the mechanical hip model will be used to further analyse commercially available hip protectors and to develop improved designs and materials for new hip protectors. Finite element modelling of hip protectors under impact loads is expected to be a valuable addition to the experimental investigations. For this purpose, a finite element model which has been developed to study design aspects of the experimental test device [63,64] will be extended in future work. In previous studies, hip protectors have been primarily assessed by the force acting directly on the greater trochanter. This force can also be determined from the force components measured in the femoral neck when using our hip model. As the force in the femoral neck is more sensitive to variations in protector position and geometry, however, femoral neck forces are considered more appropriate to assess the effectiveness of hip protectors. In a succeeding study, a new assessment criterion will be developed, in which stresses in the femoral neck are estimated from the measured forces.

Acknowledgements We thank N. Mattle and H.U. Hertach for helpful discussions and the design of the mechanical hip model and R. Bivetti and G. K¨uffer for conducting the laboratory experiments. The Swiss Council for Accident Prevention (bfu) is gratefully acknowledged for a significant portion of the project funding.

484

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485

References [1] Cummings SR, Melton LJ. Epidemiology and outcomes of osteoporotic fractures. Lancet 2002;359(9319):1761–7. [2] Harada A, Mizuno M, Takemura M, Tokuda H, Okuizumi H, Niino N. Hip fracture prevention trial using hip protectors in Japanese nursing homes. Osteoporos Int 2001;12(3):215–21. [3] Lajoie Y, Gallagher SP. Predicting falls within the elderly community: comparison of postural sway, reaction time, the Berg balance scale and the Activities-specific Balance Confidence (ABC) scale for comparing fallers and non-fallers. Arch Gerontol Geriatr 2004;38:11–26. [4] Parkkari J. Hip fractures in the elderly. Tampere: Medical School, University of Tampere; 1997. p. 85. [5] Reginster J-Y, Gillet P, Gosset C. Secular increase in the incidence of hip fractures in Belgium between 1984 and 1996: need for a concerted public health strategy. Bull World Health Organ 2001; 79(10):942–6. [6] Marks R, Allegrante JP, MacKenzie CR, Lane JM. Hip fractures among the elderly: causes, consequences and control. Ageing Res Rev 2003;2(1):57–93. [7] Hubacher M, Wettstein A. Die Wirksamkeit des H¨uftprotektors zur Vermeidung von sturzbedingten Schenkelhalsfrakturen. Bern: Schweizerische Beratungsstelle f¨ur Unfallverh¨utung, bfu; 2000. p. 109. [8] Cummings SR, Nevitt MC. Nonskeletal determinants of fractures— the potential importance of the mechanics of falls. Osteoporos Int 1994;4:67–70. [9] Greenspan SL, Myers ER, Maitland LA, Kido TH, Krasnow MB, Hayes WC. Trochanteric bone-mineral density is associated with type of hip fracture in the elderly. J Bone Miner Res 1994;9(12):1889– 94. [10] Greenspan SL, Myers ER, Maitland LA, Resnick NM, Hayes WC. Fall severity and bone-mineral density as risk-factors for hip fracture in ambulatory elderly. J Am Med Assoc 1994;271(2):128– 33. [11] Greenspan SL, Myers ER, Kiel DP, Parker RA, Hayes WC, Resnick NM. Fall direction, bone mineral density, and function: risk factors for hip fracture in frail nursing home elderly. Am J Med 1998;104(6):539–45. [12] Merilainen S, Nevalainen T, Luukinen H, Jalovaara P. Risk factors for cervical and trochanteric hip fracture during a fall on the hip. Scand J Prim Health Care 2002;20(3):188–92. [13] Norton R, Campbell AJ, LeeJoe T, Robinson E, Butler M. Circumstances of falls resulting in hip fractures among older people. J Am Geriatr Soc 1997;45(9):1108–12. [14] Parkkari J, Kannus P, Palvanen M, Natri A, Vainio J, Aho H, et al. Majority of hip fractures occur as a result of a fall and impact on the greater trochanter of the femur: a prospective controlled hip fracture study with 206 consecutive patients. Calcif Tissue Int 1999;65(3):183–7. [15] Schwartz AV, Kelsey JL, Sidney S, Grisso JA. Characteristics of falls and risk of hip fracture in elderly men. Osteoporos Int 1998;8(3):240–6. [16] Wei TS, Hu CH, Wang SH, Hwang KL. Fall characterictics, functional mobility and bone mineral density as risk factors of hip fracture in the community-dwelling ambulatory elderly. Osteoporos Int 2001;12(12):1050–5. [17] Kannus P, Parkkari J, Poutala J. Comparison of force attenuation properties of four different hip protectors under simulated falling conditions in the elderly: an in vitro biomechanical study. Bone 1999;25(2):229–35. [18] Wortberg W. H¨uft-Fraktur-Bandage zur Verhinderung von Oberschenkelhalsbr¨uchen bei a¨ lteren Menschen. Der Oberschenkelhalsbruch, ein biomechanisches Problem. Z Gerontol 1988;21:169–73. [19] Charpentier PJ. A hip protector based on airbag technology. Bone 1996;18(1):117S.

[20] Robinovitch SN, Hayes WC, McMahon TA. Energy-shunting hip padding system attenuates femoral impact force in a simulated fall. J Biomech Eng Trans Asme 1995;117(4):409–13. [21] Lauritzen JB, Petersen MM, Lund B. Effect of external hip protectors on hip-fractures. Lancet 1993;341(8836):11–3. [22] Ekman A, Mallmin H, Michaelsson N, Ljunghall S. External hip protectors to prevent osteoporotic hip fractures. Lancet 1997;350(9077):563–4. [23] Kannus P, Parkkari J, Niemi S, Pasanen M, Palvanen M, Jarvinen M, et al. Prevention of hip fracture in elderly people with use of a hip protector. N Engl J Med 2000;343(21):1506–13. [24] Meyer G, Warnke A, Bender R, Muhlhauser I. Effect on hip fractures of increased use of hip protectors in nursing homes: cluster randomised controlled trial. Br Med J 2003;326(7380):76A–8A. [25] Woo J, Sum C, Yiu HH, Ip K, Chung L, Ho L. Efficacy of a specially designed hip protector for hip fracture prevention and compliance with use in elderly Hong Kong Chinese. Clin Rehabil 2003;17(2):203–5. [26] Forsen L, Arstad C, Sandvig S, Schuller A, Roed U, Sogaard AJ. Prevention of hip fracture by external hip protectors: an intervention in 17 nursing homes in two municipalities in Norway. Scand J Public Health 2003;31(4):261–6. [27] Hubacher M, Wettstein A. Acceptance of hip protectors for hip fracture prevention in nursing homes. Osteoporos Int 2001;12(9):794–9. [28] van Schoor NM, Deville WL, Bouter LM, Lips P. Acceptance and compliance with external hip protectors: a systematic review of the literature. Osteoporos Int 2002;13(12):917–24. [29] Cameron ID, Venman J, Kurrle SE, Lockwood K, Birks C, Cumming RG, et al. Hip protectors in aged-care facilities: a randomized trial of use by individual higher-risk residents. Age Ageing 2001;30(6):477–81. [30] van Schoor NM, Smit JH, Twisk JWR, Bouter LM, Lips P. Prevention of hip fractures by external hip protectors—a randomized controlled trial. J Am Med Assoc 2003;289(15):1957–62. [31] Deprez X, Fardellone P. Nonpharmalogical prevention of osteoporotic fractures. Joint Bone Spine 2003;70:448–57. [32] Parkkari J, Kannus P, Heikkila J, Poutala J, Sievanen H, Vuori I. Energy-shunting external hip protector attenuates the peak femoral impact force below the theoretical fracture threshold—an in-vitro biomechanical study under falling conditions of the elderly. J Bone Miner Res 1995;10(10):1437–42. [33] Mills NJ. The biomechanics of hip protectors. I Mech E, part H. J Eng Med 1996;210:259–66. [34] Nabhani F, Bamford J. Mechanical testing of hip protectors. J Mater Process Technol 2002;124(3):311–8. [35] Cameron ID, Kurrle SE. Hip protectors. Lancet 2003;362:1940–1. [36] Lefauveau P, Fardellone P. Extraskeletal risk factors for fractures of the proximal femur. Joint Bone Spine 2004;71(1):14–7. [37] Viceconti M, Casali M, Massari B, Cristofolini L, Bassini S, Toni A. The ‘standardized femur program’ proposal for a reference geometry to be used for the creation of finite element models of femur. J Biomech 1996;29(9):1241. [38] Mall G, Graw M, Gehring KD, Hubig M. Determination of sex from femora. Forensic Sci Int 2000;113(1–3):315–21. [39] Courtney AC, Wachtel EF, Myers ER, Hayes WC. Effects of loading rate on strength of the proximal femur. Calcif Tissue Int 1994;55(1):53–8. [40] Pinilla TP, Boardman KC, Bouxsein ML, Myers ER, Hayes WC. Impact direction from a fall influences the failure load of the proximal femur as much as age-related bone loss. Calcif Tissue Int 1996;58(4):231–5. [41] Courtney AC, Wachtel EF, Myers ER, Hayes WC. Age-related reductions in the strength of the femur tested in a fall-loading configuration. J Bone Joint Surg Am 1995;77A(3):387–95. [42] Yang JK, L¨ovsund P. Development and validation of a human-body mathematical model for simulation of car-pedestrian collisions.. In: IRCOBI conference proceedings. 1997. p. 133–49.

S. Derler et al. / Medical Engineering & Physics 27 (2005) 475–485 [43] Bergmann G, Graichen F, Rohlmann A, Verdonschot N, van Lenthe GH. Frictional heating of total hip implants. Part 2: finite element study. J Biomech 2001;34(4):429–35. [44] Robinovitch SN, Hayes WC, McMahon TA. Prediction of femoral impact forces in falls on the hip. J Biomech Eng Trans Asme 1991;113(4):366–74. [45] Anderson MS, http://www.the3dstudio.com/models.aspx?id=95/; 1999. [46] J¨urgens HW. K¨orpermasse. In: Handbuch der ergonomie. M¨unchen: Hanser; 1989. [47] Robinovitch SN, McMahon TA, Hayes WC. Force attenuation in trochanteric soft tissues during impact from a fall. J Orthop Res 1995;13(6):956–62. [48] Parkkari J, Kannus P, Poutala J, Vuori I. Force attenuation properties of various trochanteric padding materials under typical falling conditions of the elderly. J Bone Miner Res 1994;9(9):1391–6. [49] Lauritzen JB, Askegaard V. Protection against hip fractures by energy absorption. Dan Med Bull 1992;39:91–3. [50] Askegaard V, Lauritzen JB. Load on the hip in a stiff sideways fall. Eur J Exp Muscoloskel Res 1995;4:111–6. [51] Hayes WC, Myers ER, Robinovitch SN, VandenKroonenberg A, Courtney AC, McMahon TA. Etiology and prevention of age-related hip fractures. Bone 1996;18(1):S77–86. [52] VandenKroonenberg A, Hayes WC, McMahon TA. Hip impact velocities and body configurations for voluntary falls from standing height. J Biomech 1996;29(6):807–11. [53] Cordey J, Schneider M, Buhler M. The epidemiology of fractures of the proximal femur. Inj Int J Care Inj 2000;31:56–61. [54] Ford CM, Keaveny TM, Hayes WC. The effect of impact direction on the structural capacity of the proximal femur during falls. J Bone Miner Res 1996;11(3):377–83.

485

[55] Sabick MB, Hay JG, Goel VK, Banks SA. Active responses decrease impact forces at the hip and shoulder in falls to the side. J Biomech 1999;32(9):993–8. [56] Parkkari J, Kannus P, Heikkila J, Poutala J, Heinonen A, Sievanen H, et al. Impact experiments of an external hip protector in young volunteers. Calcif Tissue Int 1997;60(4):354–7. [57] Beason DP, Dakin GJ, Lopez RR, Alonso JE, Bandak FA, Eberhardt AW. Bone mineral density correlates with fracture load in experimental side impacts of the pelvis. J Biomech 2003;36(2):219– 27. [58] Weber TG, Yang KH, Woo R, Fitzgerald RH. Proximal femur strength: correlation of the rate of loading and bone mineral density. Adv Bioeng ASME 1992;22:111–4. [59] Lotz JC, Hayes WC. The use of quantitative computed-tomography to estimate risk of fracture of the hip from falls. J Bone Joint Surg Am 1990;72A(5):689–700. [60] Lauritzen JB, Hindso K. Prevention of hip fractures with hip protectors. Orthop Int Ed 1997;5(2):125–30. [61] Lauritzen JB. Hip fractures: Incidence, risk factors, energy absorption, and prevention. Bone 1996;18(1):S65–75. [62] Wiener SL, Andersson GBJ, Nyhus LM, Czech J. Force reduction by an external hip protector on the human hip after falls. Clin Orthop Relat Res 2002;(398):157–68. [63] Schmitt K-U, Spierings AB, Derler S. A finite element approach and experiments to assess the effectiveness of hip protectors. Technol Health Care 2004;12:43–9. [64] Derler S, Spierings AB. Wirksamkeit von H¨uftprotektoren: Entwicklung eines mechanischen H¨uftmodells und eines Bewertungskriteriums. St. Gallen: EMPA Report No. 262; 2004. p. 82.