How does the way a weight is carried affect spinal loads?

People often have to carry a weight which increases the spinal load. Few in vivo measured spinal loading data exist for carrying a weight. The aim of this study was to measure the force increase on a vertebral body replacement (VBR) caused by carrying weights in different ways. A telemeterised VBR allowing the measurement of six load components was implanted in five patients suffering from lumbar vertebral body fractures. The patients carried different weights laterally in one or both hands, in front of the body and in a backpack. The force increase with respect to standing was more than twice as high for carrying a weight in front of the body compared with carrying it laterally. A weight of 10 kg in a backpack led to an average force increase of only 35 N. The position of the carried weight relative to the spine strongly affected the spinal load. PRACTITIONER SUMMARY: Carrying weights increases spinal loads. The loads on a telemeterised VBR were measured in five patients carrying weights in different ways. Holding a weight in front of the body strongly increased the force, while carrying it in a backpack led to only a minor load increase.

Loads on a vertebral body replacement during locomotion measured in vivo.

Walking is one of the most important activities in daily life, and walking exposes the spine to a high number of loading cycles. Little is known about the spinal loads during walking. Telemeterized spinal implants can provide data about their loading during different activities. The aim of this study was to measure the loads on a vertebral body replacement (VBR) during level and staircase walking and to determine the effects of walking speed and using walking aids. Telemeterized VBRs were implanted in five patients suffering from compression fractures of the L1 or L3 lumbar vertebral body. The implant allows measurements of three force and three moment components. The resultant force on the VBR was measured during level and staircase walking, when walking on a treadmill at different speeds, and when using a wheeled invalid walker or crutches. On average, the resultant force on the VBR for level walking was 171% of the value for standing. This force value increased to 265% of the standing force when ascending stairs and to 225% when descending stairs. Walking speed had a strong effect on the implant force. Using a walker during ambulation on level ground reduced the force on the implant to 62% of standing forces, whereas using two crutches had only a minor effect. Walking causes much higher forces on the VBR than standing. A strong force reduction can be achieved by using a walker.

Monitoring the load on a telemeterised vertebral body replacement for a period of up to 65 months.

PURPOSE: To determine the postoperative temporal course of the forces acting on a vertebral body replacement (VBR) for two well reproducible activities. METHODS: A telemeterised VBR was implanted in five patients. It allows the measurement of six load components. Implant loads were measured in up to 28 measuring sessions for different activities, including standing and walking. RESULTS: The postoperative temporal course of the resultant implant forces measured during standing and walking was similar in each patient, but the patterns varied strongly from patient to patient. In one patient, the forces decreased in the first year and then increased in the following 4 years. In another patient, the forces increased in the first few months and then decreased. In a third patient, the forces varied only slightly in the postoperative time. In two patients, there was a strong drop of the implant force in the first two postoperative months. The force was on average approximately 100 N or 71% higher for walking than for standing. CONCLUSIONS: The strong force reduction in the first 2 months is most likely caused by implant subsidence, and the force reduction over a period of more than 6 months is most likely caused by fusion of the vertebrae adjacent to the VBR. The short-term force increase could be attributed to bone atrophy at the index level, and the long-term force increase could be attributed to an increase in the thoracic spine kyphosis angle.

In vivo measurement of shoulder joint loads during walking with crutches.

BACKGROUND: Following surgery or injury of the lower limbs, the use of walking aids like crutches can cause high loads on the shoulder joint. These loads have been calculated so far with computer models but with strongly varying results. METHODS: Shoulder joint forces and moments were measured during crutch-assisted walking with complete and partial unloading of the lower limbs. Using telemeterized implants in 6 subjects axillary crutches and forearm crutches were compared. A force direction was more in the direction of the long humeral axis, and slightly lower forces were assumed using axillary crutches. Similar force magnitudes as those experienced during previously measured wheelchair weight relief tasks were expected for complete unloading. The friction-induced moment was hypothesized to act mainly around the medio-lateral axis during the swing phase of the body. FINDINGS: Maximum loads of up 170% of the bodyweight and 0.8% of the bodyweight times meter were measured with large variations among the patients. Higher forces were found in most of the patients using forearm crutches. The hypothesized predominant moment around the medio-lateral axis was only found in some patients. More often, the other two moments had larger magnitudes with the highest values in female patients. The assumed different load direction could only be found during partial unloading. INTERPRETATION: In general the force magnitudes were in the range of activities of daily living. However, the number of repetitions during long-lasting crutch use could lead to shoulder problems as a long-term consequence. The slightly lower forces with axillary crutches could be caused by loads acting directly from the crutch on the scapula, thus bypassing the glenohumeral joint. The higher bending moments in the female patients could be a sign of lacking muscle strength for centring the humeral head on the glenoid.

Spinal loads during position changes.

BACKGROUND: Recommendations exist how patients should change from one body position to another in order to keep the spinal loads low. However, until now it is not clear whether the loads are in fact lower if the patients follow these recommendations. The aim was to measure the loads while changing the body position. METHODS: Telemeterized vertebral body replacements have been inserted into 5 patients who had a severe compression fracture of a lumbar vertebral body. The acting loads were measured during a changing of the body position while lying and when moving from lying to sitting, from sitting to standing and vice versa. FINDINGS: When the lying patients changed their position according to the physiotherapist's recommendations, the resultant force was nearly as high as it was during relaxed standing. Otherwise, the force was nearly twice as high. Changing from a lateral lying position to sitting and vice versa caused forces of about 180% of those seen for standing when the recommendations were heeded. Without instructions, the loads were about 70% higher. Use of a trapeze bar mounted to the bed did not increase the loads. Rising from a chair with the arms hanging down laterally led to average resultant forces of 380% related to standing. Placing the hands on armrests reduced this value to 180%. INTERPRETATION: High forces may act on the spine when changing from one body position to another. These loads can be minimized when following the physiotherapist's instructions and when supporting the upper body by the arms.

An EMG-driven musculoskeletal model of the shoulder.

This paper aims to develop an EMG-driven model of the shoulder that can consider possible muscle co-contractions. A musculoskeletal shoulder model (the original model) is modified such that measured EMGs can be used as model-inputs (the EMG-driven model). The model is validated by using the in-vivo measured glenohumeral-joint reaction forces (GH-JRFs). Three patients carrying instrumented hemi-arthroplasty were asked to perform arm abduction and forward-flexion up to maximum possible elevation, during which motion data, EMG, and in-vivo GH-JRF were measured. The measured EMGs were normalized and together with analyzed motions served as model inputs to estimate the GH-JRF. All possible combinations of input EMGs ranging from a single signal to all EMG signals together were tested. The 'best solution' was defined as the combination of EMGs which yielded the closest match between the model and the experiments. Two types of inconsistencies between the original model and the measurements were observed including a general GH-JRF underestimation and a GH-JRF drop above 90° elevation. Both inconsistencies appeared to be related to co-contraction since inclusion of EMGs could significantly (p<.05) improve the predicted GH-JRF (up to 45%). The developed model has shown the potential to successfully take the existent muscle co-contractions of patients into account.

Measurement of shoulder joint loads during wheelchair propulsion measured in vivo.

BACKGROUND: Recent in vivo measurements show that the loads acting in the glenohumeral joint are high even during activities of daily living. Wheelchair users are frequently affected by shoulder problems. With previous musculoskeletal shoulder models, shoulder joint loading was mostly calculated during well-defined activities like forward flexion or abduction. For complex movements of everyday living or wheelchair propulsion, the reported loads vary considerably. METHODS: Shoulder joint forces and moments were measured with telemeterized implants in 6 subjects. Data were captured on a treadmill at defined speeds and inclinations. Additional measurements were taken in 1 subject when lifting the body from the wheelchair, using his arms only, and in 2 subjects when rapidly accelerating and stopping the wheelchair. The influence of the floor material on shoulder joint loading was accessed in 2 subjects. In general, the maximum shoulder loads did not exceed those during daily living but the time courses and magnitudes of the loads intra-individually varied much. FINDINGS: The highest forces acted during maximum acceleration and lifting from the wheelchair (128% and 188% of body weight). Grass was the only surface which led to a general load increase, compared to a smooth floor. INTERPRETATION: The increased incidence of overuse injuries in wheelchair users are probably not caused by excessive load magnitudes during regular propulsion. The high number of repetitions is assumed to be more decisive.

In vivo gleno-humeral joint loads during forward flexion and abduction.

To improve design and preclinical test scenarios of shoulder joint implants as well as computer-based musculoskeletal models, a precise knowledge of realistic loads acting in vivo is necessary. Such data are also helpful to optimize physiotherapy after joint replacement and fractures. This is the first study that presents forces and moments measured in vivo in the gleno-humeral joint of 6 patients during forward flexion and abduction of the straight arm. The peak forces and, even more, the maximum moments varied inter-individually to a considerable extent. Forces of up to 238%BW (percent of body weight) and moments up to 1.74%BWm were determined. For elevation angles of less than 90° the forces agreed with many previous model-based calculations. At higher elevation angles, however, the measured loads still rose in contrast to the analytical results. When the exercises were performed at a higher speed, the peak forces decreased. The force directions relative to the humerus remained quite constant throughout the whole motion. Large moments in the joint indicate that friction in shoulder implants is high if the glenoid is not replaced. A friction coefficient of 0.1-0.2 seems to be realistic in these cases.

Validation of the Delft Shoulder and Elbow Model using in-vivo glenohumeral joint contact forces.

The Delft Shoulder and Elbow Model (DSEM), a large-scale musculoskeletal model, is used for the estimation of muscle and joint reaction forces in the shoulder and elbow complex. Although the model has been qualitatively verified using EMG-signals, quantitative validation has until recently not been feasible. The development of an instrumented shoulder endoprosthesis has now made this possible. To this end, motion data, EMG-signals, external forces, and in-vivo glenohumeral joint reaction forces (GH-JRF) were recorded for two patients with an instrumented shoulder hemi-arthroplasty, during dynamic tasks (including abduction and anteflexion) and force tasks with the arm held in a static position. Motions and external forces served as the model inputs to estimate the GH-JRF. In the modeling process, the effect of two different (stress and energy) optimization cost functions and uniform size and mass scaling were evaluated. The model-estimated GH-JRF followed the in-vivo measured force for dynamic tasks up to about 90° arm elevations, but generally underestimates the peak forces up to 31%; whereas a different behavior (ascending measured but descending estimated force) was found for angles above 90°. For the force tasks the model generally overestimated the peak GH-JRF for most directions (on average up to 34%). Applying the energy cost function improved model predictions for the dynamic anteflexion task (up to 9%) and for the force task (on average up to 23%). Scaling also led to improvement of the model predictions during the dynamic tasks (up to 26%), but had a negligible effect (<2%) on the force task results. Although results indicated a reasonable compatibility between model and measured data, adjustments will be necessary to individualize the generic model with the patient-specific characteristics.

Realistic loads for testing hip implants.

The aim here was to define realistic load conditions for hip implants, based on in vivo contact force measurements, and to see whether current ISO standards indeed simulate real loads. The load scenarios obtained are based on in vivo hip contact forces measured in 4 patients during different activities and on activity records from 31 patients. The load scenarios can be adapted to various test purposes by applying average or high peak loads, high-impact activities or additional low-impact activities, and by simulating normal or very active patients. The most strenuous activities are walking (average peak forces 1800 N, high peak forces 3900 N), going up stairs (average peak forces 1900 N, high peak forces 4200 N) and stumbling (high peak forces 11,000 N). Torsional moments are 50% higher for going up stairs than for walking. Ten million loading cycles simulate an implantation time of 3.9 years in active patients. The in vitro fatigue properties of cementless implant fixations are exceeded during stumbling. At least for heavyweight and very active subjects, the real load conditions are more critical than those defined by the ISO standards for fatigue tests.

Loading of the knee joint during activities of daily living measured in vivo in five subjects.

Detailed knowledge about loading of the knee joint is essential for preclinical testing of implants, validation of musculoskeletal models and biomechanical understanding of the knee joint. The contact forces and moments acting on the tibial component were therefore measured in 5 subjects in vivo by an instrumented knee implant during various activities of daily living. Average peak resultant forces, in percent of body weight, were highest during stair descending (346% BW), followed by stair ascending (316% BW), level walking (261% BW), one legged stance (259% BW), knee bending (253% BW), standing up (246% BW), sitting down (225% BW) and two legged stance (107% BW). Peak shear forces were about 10-20 times smaller than the axial force. Resultant forces acted almost vertically on the tibial plateau even during high flexion. Highest moments acted in the frontal plane with a typical peak to peak range -2.91% BWm (adduction moment) to 1.61% BWm (abduction moment) throughout all activities. Peak flexion/extension moments ranged between -0.44% BWm (extension moment) and 3.16% BWm (flexion moment). Peak external/internal torques lay between -1.1% BWm (internal torque) and 0.53% BWm (external torque). The knee joint is highly loaded during daily life. In general, resultant contact forces during dynamic activities were lower than the ones predicted by many mathematical models, but lay in a similar range as measured in vivo by others. Some of the observed load components were much higher than those currently applied when testing knee implants.

In vivo measurement of shoulder joint loads during activities of daily living.

Until recently the contact loads acting in the glenohumeral joint have been calculated using musculoskeletal models or measured in vitro. Now, contact forces and moments are measured in vivo using telemeterized shoulder implants. Mean total contact forces from four patients during eight activities of daily living are reported here. Lifting a coffee pot (1.5kg) with straight arm caused an average force of 105.0%BW (%body weight) (range: 90-124.6%BW), while setting down the coffee pot in the same position led to higher forces of 122.9%BW on the average (105.3-153.4%BW). The highest joint contact forces were measured when the straight arm was abducted or elevated by 90 degrees or more, with a weight in the hand. Lifting up 2kg from a board up to head height caused a contact force of 98.3%BW (93-103.6%BW); again, setting it down on the board led to higher forces of 131.5%BW (118.8-144.1%BW). In contrast to previously calculated high loads, the contact force during passive holding of a 10kg weight laterally was only 12.3%BW (9.2-17.9%BW), but when lifting it up to belt height it increased to 91.5%BW (87-95%BW). The moments transferred inside the joint at our patients varied much more than did the forces both inter and intra-individually. Our data suggest that patients with shoulder problems or during the first post-operative weeks after shoulder fractures or joint replacements should avoid certain activities encountered during daily living e.g. lifting or holding a weight with an outstretched arm. Some energy-related optimization criteria used in the literature for analytical musculoskeletal shoulder models must now be reconsidered.

An instrumented implant for in vivo measurement of contact forces and contact moments in the shoulder joint.

To improve implant design, fixation and preclinical testing, implant manufacturers depend on realistic data of loads acting on the shoulder joint. Furthermore, these data can help to optimize physiotherapeutic treatment and to advise patients in their everyday living conditions. Calculated shoulder joint loads vary extremely among different authors [Anglin C, Wyss UP, Pichora DR. Glenohumeral contact forces. Proc Inst Mech Eng [H] 2000;214:637-44]. Additionally the moments acting in the joint caused by friction or incongruent articular surfaces, for example, are not implemented in most models. An instrumented shoulder joint implant was developed to measure the contact forces and the contact moments acting in the glenohumeral joint. This article provides a detailed description of the implant, containing a nine-channel telemetry unit, six load sensors and an inductive power supply, all hermetically sealed inside the implant. The instrumented implant is based on a clinically proven BIOMET Biomodular shoulder replacement and was calibrated before implantation by using complex mathematical calculation routines in order to achieve an average measuring precision of approximately 2%.

Design and calibration of load sensing orthopaedic implants.

Contact forces and moments act on orthopaedic implants such as joint replacements. The three forces and three moment components can be measured by six internal strain gauges and wireless telemetric data transmission. The accuracy of instrumented implants is restricted by their small size, varying modes of load transfer, and the accuracy of calibration. Aims of this study were to test with finite element studies design features to improve the accuracy, to develop simple but accurate calibration arrangements, and to select the best mathematical method for calculating the calibration constants. Several instrumented implants, and commercial and test transducers were calibrated using different loading setups and mathematical methods. It was found that the arrangement of flexible elements such as bellows or notches between the areas of load transfer and the central sensor locations is most effective to improve the accuracy. Increasing the rigidity of the implant areas, which are fixed in bones or articulate against joint surfaces, is less effective. Simple but accurate calibration of the six force and moment components can be achieved by applying eccentric forces instead of central forces and pure moments. Three different methods for calculating the measuring constants proved to be equally well suited. Employing these improvements makes it possible to keep the average measuring errors of many instrumented implants below 1-2% of the calibration ranges, including cross talk. Additional errors caused by noise of the transmitted signals can be reduced by filtering if this is permitted by the sampling rate and the required frequency content of the loads.

[Loads acting on orthopaedic implants. Measurements and practical applications]. / Die Belastung orthopädischer Implantate. Messungen und praktische Anwendungen.

The loads measured at instrumented joint replacements and other orthopaedic implants allow the optimization of their stability, wear properties, fixation stability and kinematic properties prior to clinical applications. The data obtained also indicate which activities cause very high loads and should be avoided by the patients in order not to endanger the long-term success of the implant. In addition, physiotherapy after joint arthroplasty and fractures can be further improved on the basis of these data. The technical principles for such measurements are summarized and examples for the design of load measuring instrumented implants are presented. The most important results are presented based on the measurements taken at the hip and shoulder joints, internal spinal fixation devices, vertebral body replacements and knee joints. Using this data, many practical conclusions are drawn. Due to the huge amount of data obtained from the hip, most practical advise can be provided to patients with replacement or disorders involving this joint.

In vivo glenohumeral contact forces--measurements in the first patient 7 months postoperatively.

Knowledge of forces in the glenohumeral joint is essential for understanding normal and pathologic shoulder function. It forms the basis for performing fracture treatment or joint replacement surgery, for optimizing implant design and fixation and for improving and verifying analytical biomechanical models of the shoulder. An instrumented shoulder implant with telemetric data transmission was developed to measure six components of joint contact forces and moments. A patient with humeral head arthrosis achieved good joint function after its implantation. During the first 7 postoperative months, the contact force remained below 100% BW (percent body weight) for most activities of daily living. It ranged up to 130% BW for arm positions close to the limits of motion or when acting against external resistance. When the patient tried to turn a blocked steering wheel with maximum effort, the force rose to about 150% BW, the highest level observed thus far. Of great interest were the force directions relative to the humerus, especially those in the sagittal plane, which were not greatly influenced by the type of exercise, the arm position or the external resistance. The moments due to friction in the joint reached 5.2 Nm. The friction-induced shift of contact forces relative to the implant head centre ranged up to 6.3mm. These first worldwide in vivo measurements of glenohumeral contact forces are being continued in more patients and for longer postoperative times.

Hip joint contact forces during stumbling.

AIM: To determine whether load directions for stumbling are similar to those for common activities and whether stumbling can be realistically simulated under laboratory conditions without endangering the patients. METHOD: The magnitudes and directions of hip contact forces were measured during real and simulated stumbling and compared with those found during various other everyday activities. Measurements were obtained by use of hip implants with built-in load sensors and telemetry. RESULTS: Peak forces are approximately twice as high during real stumbling as during any other activity and may range higher than eight-times the body weight. Simulated stumbling leads to much lower contact forces, especially if this happens after a warning. Accidental stumbling in everyday situations should, therefore, be avoided, especially in patients with hip replacements or arthrosis. CONCLUSIONS: The directions of peak hip contact forces relative to the femoral bone are nearly constant for any activity, including real stumbling. This observation supports the assumption that muscle and bone anatomy plus muscle function are optimized in order to minimize stresses in bone and muscles. Any impairment of such a mechanically balanced system will increase the musculoskeletal loads. Malposition of total hip implants or muscle deficits caused by the surgical approach must, therefore, be avoided or minimized.

ISSLS prize winner: A novel approach to determine trunk muscle forces during flexion and extension: a comparison of data from an in vitro experiment and in vivo measurements.

STUDY DESIGN: Disc pressure and fixator load were measured in an in vitro setup and compared to in vivo measurements with the identical transducers from the two groups participating in this study. OBJECTIVES: The goal of this in vitro study was to determine the magnitude of trunk muscle forces during flexion and extension. The loading conditions in this study accounted for body weight, local and global muscles, and forces resulting from the support of the abdominal soft tissue in different postures. Resulting intersegmental motions and intradiscal pressure in each segment and the six load components in both rods of an internal fixator were determined. SUMMARY OF BACKGROUND DATA: The spine is primarily stabilized by muscle forces, which greatly influence spinal loads. However, little information exists on the magnitudes of trunk muscle forces during postures like flexion and extension of the upper body. METHODS: Seven human cadaveric lumbar spines were mounted in a spine tester and adjusted to different degrees of flexion and extension of the upper body with different hip flexions. For each specimen, a total of 124 load cases were studied. They included combinations of a vertical compressive load, a follower load and forces pulling with cables at a plate fixed at the cranial end of the specimen to simulate rectus abdominis, erector spinae, and a supporting force of the abdomen. The muscle forces were varied until the external moment, necessary to keep the lumbar spine specimen in the examined posture, was zero. This was achieved with different muscle force combinations. Loads on internal fixators as well as intradiscal pressure and intersegmental rotation at all levels were measured. The muscle force combination that caused intradiscal pressures and loads in the internal fixator closest to those measured in vivo were assumed to be the muscle forces which can be expected in vivo. RESULTS: Generally, intradiscal pressure was closer to in vivo measurements than the fixator loads. The force in the m. erector spinae increased with the flexion angle but was only slightly influenced by extension. The estimated forces in the erector spinae were 100 N for standing, 130 N for 15 degrees extension, and 520 N for 30 degrees flexion of the upper body. Little influence was found on the intersegmental motion. CONCLUSION: In vitro loading conditions can be approximated closely to in vivo conditions with the simulation of an axial preload, local, and global muscles. This novel approach can help to estimate muscle forces, which can usually not be measured. The results from this study provide important input for FEM models, which may then allow the investigation of different load cases.

Direct comparison of calculated hip joint contact forces with those measured using instrumented implants. An evaluation of a three-dimensional mathematical model of the lower limb.

Characterisation of hip joint contact forces is essential for the definition of hip joint prosthesis design requirements. In vivo hip joint contact force measurements have been made using instrumented hip joint prostheses. However, to allow determination of the range of values of joint contact force and their directions relative to anatomical structures in a range of subject groups sufficient to form an agreed data base it is necessary to adopt a different approach without the use of an implanted transducer. The use of mathematical models of the lower limb to examine the forces in soft tissues and at the joints has provided valuable insight into internal loading conditions. Several authors have proposed mathematical musculo-skeletal models. However, there have been only limited attempts at validation of these models. It is possible to use the results of in vivo force measurements from instrumented prostheses to validate the results calculated using the mathematical models. In this study two subjects with instrumented hip joint prostheses were studied. Forces at the hip joints were calculated using a three-dimensional model of the leg. Walking at slow, normal and fast speeds (0.97-2.01m/s), weight transfer from two to one leg and back again, and sit to stand were studied. Direct comparisons were made between the 'gold standard' measured hip joint contact forces and the calculated forces. There was general agreement between the calculated and measured forces in both pattern and magnitude. There were, however, discrepancies. Reasons for these differences in results are discussed and possible model developments suggested.

[Spinal load bearing during sitting in an office chair with a tilting back]. / Wirbelsäulenbelastung beim Sitzen auf einem Bürostuhl mit nach hinten kippbarer Rückenlehne.

Long periods of quiet sitting is considered a cause of low back pain. It is often assumed that spinal loads are high, especially when sitting erect. Modern office chairs with a tiltable back permit changes in the seated posture. In the most reclined position, some new chairs even match a kyphotic form of the lumbar spine. It is assumed that sitting on such a chair reduces low back pain. With the aim of determining spinal loading in different sitting positions, the loads acting on implanted fixation devices were measured telemetrically in two patients. Loads were measured in patients sitting on six different chairs with tiltable backs. In modern chairs, implant loading was always lower than while walking. In the end-tilt position of the chairback, loads were always lower than when the chairback was upright. Even when the lordotic curvature of the lumbar spine was "corrected", loads on the fixator were lower than when the subject was seated in the upright position. In a modern chair, spinal loading is no higher than with non-adjustable office chairs.