Physical human-robot interaction of an active pelvis orthosis: toward ergonomic assessment of wearable robots

Five healthy adults (74.4 ± 6.8 kg, 1.73 ± 0.07 m, 29.2 ± 6.3 years old) were enrolled for the study. All participants signed an informed consent before starting the experimental sessions. The research procedures were conducted at the premises of Fondazione Don Carlo Gnocchi (Firenze, Italy) in accordance with the Declaration of Helsinki, after the approval of the local Ethical Committee.

The APO is a bilateral powered exoskeleton, and it is constituted of three main subsystems: the mechanical structure, the actuation units, and the control system. In the following, we provide a description that summarizes its main features.

The mechanical structure of the APO is symmetrical with respect to the sagittal plane (Fig. 1). Each side of the robot is composed of two main subsystems, namely, the chain of passive DoFs and the transmission means that transfer the assistive torque from the actuation unit to the human hip articulation.

Each chain of passive DoFs originates from a main posterior carbon fiber plate (i.e., the frame of the chain of passive DoFs), which connects the exoskeleton to the wearer’s trunk by means of an orthotic cuff. Another carbon-fiber plate can slide horizontally against the frame by means of a passive translational DoF (axis number 1 in Fig. 1). The sliding plate houses two passive rotational DoFs, whose axes of rotation (namely, axes number 2 and 3 in Fig. 1) are orthogonal and cross each other at point P. Thanks to the combined action of the translational passive DoF, the rotation axis number 2 can be aligned with the human hip a/a axis. The range of motion of axis number 2, and, consequently, that of the user’s a/a, is restricted to −15° to +20° via mechanical end stops. The concomitant movement of the translational passive DoF and that of axis number 3 allow the user to also have a free hip i/e. The ROM of axis number 3 is restricted to –10° to +10° by mechanical end stops. This kinematic chain of passive DoFs connects the main carbon fiber plate to a lateral arm, also made of carbon fiber. The distance between the left and right lateral arms on the frontal plane can be manually adjusted to fit the width of the wearer’s pelvis. Each lateral arm is made of two telescopic shells that can slide against each other in order to maximally align the human and robot f/e axes in the sagittal plane by manually tuning their lengths. A thigh linkage rotates around the f/e axis and couples with the wearer’s thigh via an orthotic shell. Finally, an additional rotational DoF is inserted between each link and the cuff (axis number 4 in Fig. 1); this allows the alignment of the cuff and thigh longitudinal axes, thus providing considerable stability during movement. The APO kinematic chain design is patent pending [30].

The transmission system connects the actuation unit placed on the rear part of the lateral arm to a driven pulley placed coaxially with the hip f/e axis by means of a steel cable (U8191517, Carl Stahl®GmbH, Suessen, Germany) in a capstan configuration. The stiffness of the cable is 250 N/mm, equivalent to a torsional stiffness at the hip joint of 756 Nm/rad. The rotation axis of the actuation unit is parallel to the hip f/e axis.

Each actuation unit is a series elastic actuator (SEA). Each SEA is composed of a 70 W DC motor (EC60, Maxon Motor®, Sachseln, Switzerland), a 100:1 harmonic drive (CPL-14A-100-2A, Harmonic Drive®, Limburg, Germany), and a custom torsional spring (patent pending) having a stiffness of 100 Nm/rad. Two absolute 17-bit Rotary Electric Encoder™ units (DS-37 and DS-25 Netzer Precision Motion Sensors Ltd, Misgav, Israel) measure the spring deformation and the actual hip joint angle, respectively.

The control system has a hierarchical structure made of a low- and a high-level layer. For the control system, we adopted the same control architecture as that described in [22]; hereafter, its main features are recapped for the sake of clarity. The low-level layer is a closed-loop torque control. The controller is a 2-pole-2-zero compensator. The closed-loop compensator allows for a relatively high closed-loop bandwidth (namely, 15.5 Hz) and low joint residual parasitic stiffness (lower than 1 Nm/rad in the typical frequency range of walking). The high-level assistive control aims at computing a desired assistive torque profile during the stride for each of the two powered hip joints. It is based on the model-free algorithm presented in [31]. It relies on adaptive oscillators, mathematical tools [32] that—when coupled with a kernel-based non-linear filter—can track, estimate, and predict quasi-periodic signals (e.g., hip angles during gait) with zero-delay. Hence, during ground-level walking tasks, it is possible to determine the phase φ, frequency, and envelope of each hip joint angle and to reliably predict the joint angle during the stride period thanks to an adjustable phase shift Δφ (Δφ = 0.628 rad in this work). The assistive reference torque is provided by: \( {\uptau}_{\mathrm{des}}={\mathrm{K}}_{\mathrm{v}}\cdot \left[{\widehat{\uptheta}}_{\mathrm{j}}\left(\upvarphi +\Delta \upvarphi \right)-{\widehat{\uptheta}}_{\mathrm{j}}\left(\upvarphi \right)\right] \), where K_v [Nm/rad] is an adjustable virtual stiffness and \( \widehat{\uptheta_{\mathrm{j}}}\left(\upvarphi \right) \) and \( \widehat{\uptheta_{\mathrm{j}}}\left(\upvarphi +\Delta \upvarphi \right) \) are the hip joint angle estimate and its predicted future value, respectively. This means that thanks to the virtual stiffness, we can attract the hip f/e angle from the current position \( {\widehat{\uptheta}}_{\mathrm{j}}\left(\upvarphi \right) \) to the future \( {\widehat{\uptheta}}_{\mathrm{j}}\left(\upvarphi +\Delta \upvarphi \right) \) as a result of the application of a torque τ_des. In this experiment, the parameters of the assistive controller were set according to [31].

All volunteers walked barefoot on a treadmill at three different speeds (slow, normal, and fast, named V1, V2, and V3, respectively) and under five modalities, namely: (i) without wearing the APO mechanics, but with the pelvis orthotic cuff (natural walking, NW); (ii) wearing the APO in the zero-torque control mode (transparent mode, TM); (iii)–(v) assistive mode with three different levels of assistance (low, moderate, and high assistive modes, named AM1, AM2, and AM3, respectively). The velocity V2 was selected according to the principle of dynamic similarities [33] and was thus calculated as \( \mathrm{V}2=\sqrt{{\mathrm{F}}_{\mathrm{r}}\cdot \mathrm{g}\cdot \mathrm{L}} \), where F_r is the Froude number, g is the gravity constant, and L is the leg length (measured from the greater trochanter prominence to the lateral malleolus). In this experiment, F_r = 0.1. V1 and V3 were selected to be equal to V2 ± 0.25 V2. Each subject walked in all conditions (in the order: NW, TM, AM1, AM2, and AM3) at all different speeds (in the order: V1, V2, and V3). Each trial consisted of 20 strides.

The desired assistance level was set according to the following methodology. During a familiarization session, each volunteer was requested to walk at V3 while the experimenter progressively increased the value of K_v. The value of K_v for AM3 was that corresponding to the highest level of assistance that the subject considered as comfortable. In fact, all subjects reported discomfort for high values of K_v. When the peak torque—normalized to body weight—exceeded an across-subjects average value of 0.14 Nm/kg, the human and robotic hip joint kinematics difference increased, thus resulting in an assistive action that is not compliant with human biomechanics. We will further discuss this issue in the discussion session. Once the K_v upper limit was identified, it was scaled down by 33 and 66% for the AM2 and AM1 conditions, respectively.

Owing to the APO modular architecture, it was possible to wear only the pelvis cuff in NW. In all other assisted conditions, including TM, thigh cuffs were also worn.

The APO actual hip joint angle and torque were measured by using the information from the encoders. Lower-limb kinematics and the movement between the orthotic cuffs and the corresponding body segments were measured by means of an optoelectronic system (SmartD, BTS, Milan, Italy) detecting spherical passive markers placed on specific points on the robot and user. Standard software (Smart Tracker, BTS, Milan, Italy) was used to compute the 3D coordinates of the markers. Ad-hoc post-processing was performed in the MATLAB environment (MathWorks, Natick, MA, USA). Human kinematics was calculated according to an adapted LAMB model [34] (see Fig. 2), in which the required posterior superior iliac spine landmark, hidden by the exoskeleton, was reconstructed through a marker placed on top of a pelvis-anchored stick. All acquired data were segmented according to heel strikes and resampled from 0 to 100% of the stride cycle. The heel strikes were detected with a dedicated algorithm that processes the 3D coordinates of feet markers.

From the collected data, we computed the following variables, which provided a quantitative assessment of the level of ergonomics: (i) the root mean square (RMS) of the difference of the hip, knee, and ankle f/e angle—measured by the motion capture system—between the NW and TM/AM conditions; we named these variables ‘human hip angle deviation’ (H-HAD), ‘human knee angle deviation’ (H-KAD), and ‘human ankle angle deviation’ (H-AAD), respectively; (ii) the RMS of the difference between the APO hip f/e angle, recorded by the joint encoder, and the anatomical hip f/e angle, computed through the motion capture system; we named this variable ‘human-robot hip angle deviation’ (HR-HAD); (iii) the standard deviation (SD) of the relative displacements between the markers placed on the orthotic cuffs of the APO and those placed on their corresponding body segments; we named this set of variables ‘physical human-robot interface displacement’ (pHR-ID); (iv) the human joints’ ROM in the sagittal plane (hip, knee, and ankle) and in the frontal plane (only at the hip); and (v) spatio-temporal parameters (i.e., step length, stance time, and cadence).

In the case of the pHR-ID calculation, we used the SD instead of the RMS because it is more adequate for measuring the relative displacements of two points in space, without considering the constant offset between them.

The reference value used to evaluate H-HAD, H-KAD, and H-AAD in all walking conditions was the average intra-subject variability, defined as the among-subjects average of intra-subject variability. The intra-subject variability, calculated in each walking condition for one subject, is here defined as the average SD in a stride period. Hence, the intra-subject variability quantifies the natural differences of the f/e angles for one subject, while the average intra-subject variability captures the mean of these natural kinematic differences considering all subjects.

The relative displacement between markers placed on the same rigid body, namely, the pelvis cuff of the APO during walking, averaged over all conditions, was taken as the noise level of the experimental setup (2.3 ± 0.8 mm).

Table 1 reports the spatio-temporal parameters for different trials. When considering speeds, the stance time shows a negative trend associated with speed increase, while cadence shows a positive trend; conversely, none of the spatio-temporal parameters show any trend associated with the assistance level. All spatio-temporal parameters, except step length, show significant differences due to speed increase. When comparing natural walking with the TM/AMs, all spatio-temporal parameters except the stance time show no significant differences. In particular, the average stance time in AM2 is slightly larger than in NW. When comparing walking conditions in which the exoskeleton was worn, no statistically relevant differences are found, except for the stance time, which shows a difference between TM and AM2.

In Fig. 3, we present the kinematics of the hip, knee, and ankle in the sagittal plane and the kinetics of the hip for one representative subject.

The hip and knee ROMs (see Fig. 4 and Table 2) show a positive trend associated with assistance and speed increase. With respect to NW, the hip ROM significantly increases when wearing the exoskeleton in the TM and AMs, while the knee ROM significantly increases in the AMs but not in the TM; the ankle ROMs do not show any significant difference, except for AM3. When comparing the AMs, we notice no significant differences in the ROMs for the hip and ankle, while the knee ROM in AM1 is significantly lower than those in AM2 and AM3.

Despite the hip and knee ROM increase with assistance, the overall f/e angle trajectories in all walking conditions are highly overlapped, showing consistency between the kinematics in NW and in other walking conditions. In fact, when considering H-HAD, H-KAD, and H-AAD (see Fig. 5 and Table 3), which account for the global difference between the kinematics in NW and in other conditions, we obtain values comparable with the average intra-subject variability of the hip, knee, and ankle f/e angles (see Fig. 6 and Table 4). H-HAD ranges from 1.8 ± 0.8° in V1-TM to 3.9 ± 1.1° in V2-AM3, while the average intra-subject variability of the hip f/e angle ranges from 1.5 ± 0.5° in V3-NW to 2.4 ± 1.0° in V3-AM3. H-KAD ranges from 2.6 ± 1.5° in V2-TM to 5.7 ± 1.5° in V1-AM2, while the average intra-subject variability of the knee f/e angle ranges from 1.6 ± 0.2° in V3-NW to 3.4 ± 0.6 in V1-AM2. H-AAD ranges from 1.7 ± 0.8° in V2-TM to 3.8 ± 3.0° in V1-AM2, while the average intra-subject variability of the ankle f/e angle ranges from 1.4 ± 0.2° in V3-NW to 2.2 ± 0.2 in V1-AM2.

When comparing the AMs, H-HAD and H-KAD show no statistically significant differences related to the assistance level increase. Conversely, H-HAD and H-AAD show significant differences when considering TM vs AM2/AM3.

H-HAD is not significantly affected by speed increments, while H-KAD in V1 shows significant differences from V2 and V3; H-AAD in V1 is significantly different from V3.

Table 5 and Fig. 7 report the pHR-ID data for all trials and cuffs. When considering the walking modalities, we notice a positive trend in the displacements between the cuffs and their corresponding body segments that is related to the increment of delivered assistance. When considering speeds, only thigh cuffs show a positive trend associated with speed increase. The natural walking condition is only reported for the pelvis, since thigh cuffs could not be worn without the APO being donned.

For the pelvis cuff, pHR-ID, which ranges from 2.0 ± 0.6 mm in V2-TM to 3.1 ± 2.0 mm in V3-AM2, shows non-significant differences related to speed and significant differences associated with the walking modalities.

For the right cuff, the pHR-ID values, which range from 4.0 ± 0.6 mm in V1-TM to 6.3 ± 1.2 mm in V3-AM3, show significant differences related to both speed and assistance increments.

In Fig. 8, we show the human and APO kinematics at the hip f/e joint for one representative subject. We report a positive trend in the difference between the two trajectories that is associated both with speed and assistance increment. This difference arises between 40 and 90% of the stride cycle (torque delivery intervals from all the subjects are included in this range), when the flexion torque is delivered. When comparing walking modalities, all differences in HR-HAD are significant; when comparing speeds, the differences between V3 and V1/V2 are significant. HR-HAD ranges from 2.7 ± 0.8° in V1-TM to 7.2 ± 1.1° in V3-AM3 (see Fig. 9 and Table 6).

From the analysis of the spatio-temporal parameters, we deduce that speed significantly influences cadence and stance time. These results are expected and consistent with normative data [36]. The effect of wearing the APO in the TM and AMs on the cadence and step length is not significant, while the effect on the stance time is significant but not relevant (~1% of stride cycle variation). Hence, from the point of view of the spatio-temporal parameters, the APO does not affect physiological walking either in the TM or in the AMs.

When considering the ROM, the effect of speed is expected and consistent with normative data. The effect of assistance both on the hip and knee f/e is significant, but the increased hip flexion is compensated by an increased knee flexion, which maintains the step length stable. This is most probably due to the natural dynamic coupling between the thigh and shank, which induces an increased f/e angle peak at the knee in response to the torque delivered at thigh level. This result is in line with the study reported in [31], in which the assistive controller used in this work was validated on the treadmill-based LOPES platform. This compensation mechanism seems not to involve ankle f/e.

While the ROM is an excellent indicator of the overall joint kinematics, it only accounts for a peak-to-peak difference over the stride. For this reason, it is interesting to consider parameters with higher content of information regarding the deviation from natural kinematics over the stride period. H-HAD, H-KAD, and H-AAD take into consideration a point-by-point difference of the joint angle trajectories between NW and other walking modalities during each interval of the stride cycle. A major deviation from natural kinematics would be a symptom of a hindering effect of the APO on the human joint. Nevertheless, these indicators are of the same order of magnitude as the average intra-subject variability (see Table 3 and Table 4). The deviation from natural kinematics introduced by wearing the exoskeleton in the TM/AM can be considered comparable with physiological kinematics variability and, hence, negligible. Therefore, the APO hinders physiological joint kinematics minimally. Furthermore, the deviation from NW kinematics is smaller in the TM than in the AMs for hip and ankle joints, while the differences are not significant for the knee joint (see Fig. 5). This proves that wearing the exoskeleton in the TM produces the lowest level of deviation from physiological walking among walking modalities.

It is worth mentioning that the registered average pelvis anteversion of 10 ± 1° caused by wearing the APO is consistent with previous studies, but additional research should be done in order to assess if such a deviation may cause noxious effects in terms of injury or discomfort in young adults carrying a light-weighted backpack (8 kg). Previous research has indicated an increased lumbosacral force in the case of 15%–20% of body weight carriage [37]; however, in this study, the APO represents in average 10% of the body weight of the participants.

Displacements between the cuffs and corresponding body segments are indicators of possible joint rotation axis misalignments and may produce pressure sores during exoskeleton usage. Absence of relative displacements is hardly obtainable owing to skin elasticity and compliance of soft tissue, although they must be kept as low as possible. The results show the pelvis cuff to be the most stable among all the cuffs in terms of pHR-ID. Indeed, the averaged pHR-ID value across speeds during NW is 2.2 ± 0.2 mm, while in the case of assistance, it takes its highest value of 3.2 ± 2.2 mm, only 1 mm higher. For the right thigh cuff, pHR-ID is in general two times higher than that of the pelvis cuff but still not critical in terms of pain or risk of injury for the user. In fact, in [38], the authors reported no skin damage or pain for skin strains of up to 11.7 mm on the forearm. The difference between the thigh and pelvis pHR-IDs is mainly due to the following. (i) The presence of large active muscles that can cause shifts of both cuffs and thigh markers. (ii) The different geometrical properties of the cuffs; the quasi-cylindrical shape of the thigh cuffs facilitates parallel and rotational shifts along the femoral axis. Instead, the pelvis cuff is molded to firmly lean against the iliac crests of the wearer, preventing downward shifts. Its geometry also impedes rotational shifts around the vertical axis. (iii) The inertia of moving masses, which increases more consistently at the thigh level, particularly at higher speeds, i.e., for larger accelerations. Since pHR-ID is below 1 cm, we can conclude that the APO is very unlikely to cause discomfort or injury due to excessive displacement between the cuff and body parts under normal use (i.e., when the peak torque—normalized to bodyweight—remains below 0.14 Nm/kg). This is confirmed by the fact that no subject reported pain or discomfort during usage.

One of the main limitations of this study is the lack of a direct measure of JAxM between the human and the robotic hip joints. Nevertheless, HR-HAD together with pHR-ID may provide important information regarding the appearance of possible injuries when using an exoskeleton. In the case or rigid WRs, the presence of large or small HR-HAD and pHR-ID can only imply large or small JAxM, respectively, since the human-robot kinematic configuration is uniquely determined by the human-robot kinematic discrepancy and the body-cuff relative displacements. The exoskeleton presented in this study has a certain degree of mechanical compliance owing to: (i) the finite stiffness of the structure, (ii) mechanical backlash, and (iii) the integration of pDoF mechanisms; the former two items are negligible with respect to the latter one. Despite compliance, when pHR-ID and HR-HAD are large, they can be considered as indicators of major JAxM, which may produce pain or discomfort, since JAxM is the only plausible cause of large pHR-ID and HR-HAD. Instead, when small HR-HAD and pHR-ID are measured, they cannot provide information regarding major or minor JAxM because the compliance of pDoF may absorb the parasitic forces induced by JAxM. Nevertheless, the evaluation of pHR-ID and HR-HAD may provide insight regarding eventual pain or discomfort to the user. Indeed, if natural kinematics is not altered by the exoskeleton (H-HAD, H-KAD, H-AAD) and no pain or discomfort is reported by the wearer at the physical interface (low pHR-ID) and articulations, possible JAxM may be considered not critical for the exoskeleton ergonomics in terms of risk of injury.

The presented results show a significant increment in the human-robot kinematics discrepancy associated with both speed and assistance level increase. When wearing the APO in the TM, the HR-HAD is the lowest; then, it increases with the increment of delivered assistance (namely, with the increment of the proportional gain K_v), mostly between 40% and 90% of the gait cycle. It is worth noting that since the assistive strategy relies on the APO kinematics to determine the delivered torque profile, the larger the HR-HAD, the worse the match of the delivered torque envelope with the articulation biomechanics. Furthermore, the data show that HR-HAD is significantly dependent on the level of assistance (K _v ). Thereby, exceeding a body-weight-normalized torque limit of 0.14 Nm/kg would result in both a suboptimal human-robot kinematic coupling and a torque profile that is not compliant with human biomechanics, thus perceived as uncomfortable by the wearer. Hence, the measured HR-HAD seems to be not critical in terms of comfort and ergonomics, considering that the average peak torque—normalized for bodyweight—is below 0.14 Nm/kg.

Nevertheless, possible causes of HR-HAD are: (i) compression of soft tissues at the thigh and pelvis level and shifts of the orthotic shells during torque delivery (namely, pHR-ID), (ii) compliance of the APO mechanical structure (pDoFs), and (iii) misalignment between human and APO hip joint rotation axes (JAxM). Regarding the first cause, we observe that body-relative cuff shifts coincide with pHR-ID, which is composed of two components: one perpendicular to the interaction surface (e.g., compression of soft tissues) and one tangential (e.g., skin stretching and friction). Even in the worst case, in which we assume pHR-ID to be equal to one of the two components (e.g., only the tangential), the effect on comfort and risk of injury is negligible [38], regardless of direction. The other two causes are closely related since, in the case of JAxM, the parasitic load on the hip articulation may be absorbed by the chain of pDoFs, inducing a change in the APO kinematic configuration, which will be reflected by an increased HR-HAD.

Since physiological kinematics is only minimally perturbed by the APO (low H-HAD, H-KAD, H-AAD) and no subject reported pain during the execution of all trials, as confirmed by the low pHR-ID, we may consider the APO ergonomic, regardless of actual JAxM.