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Pneumatic active gait orthosis
Mechatronics 11 (2001) 301±323
Pneumatic active gait orthosis G. Belforte, L. Gastaldi*, M. Sorli Department of Mechanics, Politecnico di Torino, c.so Duca degli Abruzzi, 24 10129 Torino, Italy Received 30 July 1999; received in revised form 8 February 2000; accepted 1 March 2000
Abstract The paper describes design, construction and experimental testing of an active gait orthosis intended to assist locomotion in paraplegic subjects. Design speci®cations were formulated on the basis of an analysis of the context in which the device will be used. Dierent joint activation solutions were then proposed and examined. Considerable care was devoted to de®ning the logic and control system and to implementing the electropneumatic circuit, given that the system must be worn by a disabled subject who has no sensory perception in the lower limbs. Experimental tests, which made it possible to determine system performance at each step of development, were carried out with no user, then with a healthy user and ®nally with a paraplegic user. Experimental results consisting of graphs and photographic images are presented and discussed. 7 2000 Elsevier Science Ltd. All rights reserved. Keywords: Pneumatic; Active orthosis; Paraplegia
1. Introduction Signi®cant resources continue to be invested in the development of orthoses for functional gait restoration in paraplegic patients [1]. There can be no doubt that being able to maintain an upright position is psychologically important. Having to use a wheelchair in surroundings which are normally designed for people who can walk, not only makes everyday life immeasurably more complicated, but also * Corresponding author. Tel.: +39-011-564-6939; fax: +39-011-564-6999. E-mail address: [email protected] (L. Gastaldi). 0957-4158/01/$ - see front matter 7 2000 Elsevier Science Ltd. All rights reserved. PII: S 0 9 5 7 - 4 1 5 8 ( 0 0 ) 0 0 0 1 7 - 9
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aects personal relationships and communication [2,3]. From the physiological standpoint, a number of therapists maintain Ð though without experimental documentation Ð that the upright position promotes better body system function, prevents bedsores, and combats osteoporosis, while in Ref. [4], the physiological bene®ts of an upright position, at least in the short term, are claimed to be not so great. In approaching this problem, however, it must be borne in mind that restoring locomotion is not generally seen by the disabled as a fundamental need, and in any case depends on how it is accomplished. After the initial enthusiasm at being able to stand and walk again, in fact, the patient is faced with serious psychological and physiological problems, involving the cardio-vascular, the urological and the respiratory apparatus, in comparison with which walking becomes a secondary consideration. In addition, the patient must cope with the diculties associated with locomotion itself and with the orthopaedic aids, including high energy expenditure, diculty in putting on the orthosis, poor cosmesis, and unnatural gait [5]. All this leads to a feeling of disappointment, which in most cases at least [3,6], causes the patient to abandon the orthosis after initial rehabilitation at the orthopaedic clinic. Thus, the functional need to make the upright position part of the disabled subject's daily life once again is oset by an insucient patient motivation, which is partly justi®ed by the fact that the wheelchair is an excellent means of covering considerable distances with a modest energy expenditure comparable to that of normal gait [7]. Currently, the only means whereby the upright position and locomotion can be recovered are passive orthoses. These devices make up for the inability of paralysed postural muscles to contract by stiening the lower limbs. The passive orthoses which provide best performance are the reciprocating types, which enable the user to assume an upright posture without stabilising aids and can also be used by patients with high-level lesions. Walking, on the other hand, calls for the use of crutches or a walking frame and, as it takes place with the legs fully extended, is extremely tiring [8,9]. Other prototype systems involve the use of functional electrical stimulation (FES), but the large number of channels and the resulting muscle fatigue limit the distance that can be covered. Hybrid systems requiring fewer channels are more promising, as FES is associated with a support exoskeleton [10]. Another approach is that of using active orthoses, i.e., orthoses in which the energy required for locomotion is provided by an external energy source. A prototype system of this kind was constructed by Vukobratovic [11]; here, system stabilization and the fact that all lower limb joints are activated in the sagittal plane makes this structure heavy, bulky and complex from the control standpoint. A more recent investigation of walking machines was conducted by Rabischong et al. [12,13] who proposed a master slave control system for training paraplegic patients to walk again. This system consists of a passive orthosis worn by the physiotherapist and an active orthosis worn by the patient, which is equipped with actuators so that the trainer's movements are faithfully reproduced. A third active
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orthosis was proposed in Refs. [14,15] by Miyamoto et al. Hip and knee joints are activated hydraulically and equilibrium control is maintained by the patient. Generally, these solutions cannot yet be regarded as viable alternatives to the wheelchair. They are used for therapeutic purposes in order to achieve long-term bene®ts, or for mobility in structured environments such as the home, workplace or school. Higher walking speeds and lower eort levels are two requisites which would clearly encourage more extensive orthosis use, both by those who already wear an orthosis on a regular basis, and by those who do not use their orthosis because of the excessive energy expenditure involved. The active orthosis developed at the Department of Mechanics of the Politecnico di Torino was designed to provide these requisites and, where possible, those indicated in Ref. [5] as essential, viz., cosmesis, wearability, relative simplicity, etc. As these parameters are fundamental to the structure's acceptance by paraplegic patients, the trade-o which best satis®es the user must be identi®ed. Formulation of design speci®cations was preceded by an analysis of the operation of passive orthoses (Hip Guidance Orthosis, Reciprocating Gait Orthosis, Advanced Reciprocating Gait Orthosis) and active prototypes, while consideration was also given to minimizing the energy cost of locomotion [16,17]. Throughout orthosis development, moreover, rehabilitation therapists and paraplegic patients provided valuable feedback. The prototype consists of a passive support structure and a pneumatic actuation system which powers and controls the hip and knee joints. Possible design solutions were discussed in Refs. [18,19], where initial tests carried out on the orthosis without a wearer were also presented. The results of tests conducted with healthy subjects were presented in Ref. [20]. This paper presents the ®nal version of the active orthosis prototype, the actuation and control system, and the results of experimental tests performed with a paraplegic user. 2. Gait pattern The desirable properties of an orthosis can be summed up as providing the ability to walk, with energy savings, the ability to provide near-normal gait and a highly wearable structure. It is important to evaluate which degrees of freedom must be introduced and which must be blocked in order to have satisfactory gait with the minimum number of actuators. To do so, we assessed the in¯uence of each main characteristic of reciprocating gait [16] on the trajectory of the body's center of gravity. This analysis was carried out only on the sagittal plane, as it is there that the larger movements occur. The hip, knee and ankle rotation angles in the sagittal plane resulting from analysis of physiological gait are shown in Fig. 1 [21]. Hip ¯exion permits the lower limbs to move forward in alternation; at heel
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strike, the hip is in ¯exion and the trunk is situated to the rear of the foot that has been swung forward, while in the subsequent stance phase the hip is fully extended. Knee ¯exion occurs both during the stance phase, where it minimizes the center of gravity's vertical excursion, and during the swing phase where it prevents the foot from hitting the ground. Thus, knee ¯exion essentially reduces energy consumption. Plantar ¯exion takes place at various times during gait: when the heel contacts the ground and the ankle moves downwards and to the rear, at the terminal swing phase to prevent the foot from dragging, and at the end of the stance phase when the body's entire weight must be propelled forward. From these considerations, ¯anked by an assessment of costs versus bene®ts, it follows that the hip and knee joints in the sagittal plane are indispensable for near-normal gait and for reducing the excursion of the body's center of gravity and, consequently, energy expenditure.
Fig. 1. Hip, knee and ankle joint rotations in the sagittal plane.
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Stick diagrams for evaluating the gait pattern that can be achieved with the assumptions made above are shown in Fig. 2 for the various phases, both in the sagittal and transversal plane. The balance is maintained by the user with a pair of crutches. Diagram A in Fig. 2 shows the instant of initial contact between the right foot and the ground, which corresponds to the beginning of the stance phase and of the double support period (the hatched footprints indicate contact with the ground). The subsequent loading response phase (B) prepares for the left limb's swing motion: body weight is transferred from the left to the right limb, and the left crutch is positioned anteriorly in order to ensure balance in the subsequent phases. In the midstance phase (C), the swinging limb moves ahead of the stance limb. Forward movement is produced by ¯exion of the left hip, and by ¯exion of the knee, which prevents the foot from hitting the ground. Diagram D shows the terminal moment of the single limb stance phase, when the left knee is extended prior to contact with the ground. Diagram E showing preparation for the swing phase and diagrams F, G and H, which correspond respectively to the initial swing, midswing and terminal swing phases, are symmetrical to the previous diagrams, given that they represent the right leg's swing motion. 3. Exoskeleton The prototype's support structure is a commercial passive reciprocating gait orthosis, modi®ed to introduce new degrees of freedom and accommodate the actuators. As only two-dimensional gait was assumed, the knee and hip joints consist of cylindrical hinges and ankle rotation is prevented by immobilizing the joint with an Ankle Foot Orthosis (AFO). The choice of a commercial structure such as the ARGO (Advanced Reciprocating Gait Orthosis) [22] made it possible to use a structure which has already been optimized, particularly as regards the hip hinge joints and the rear
Fig. 2. Gait phases with the new active orthosis.
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cable. The latter rigidly connects the two hips to generate a reciprocating movement: in other words, extension of one hip empowers ¯exion of the contralateral hip and vice versa. As regards psychological acceptance, most paraplegics are already familiar with the structure. A further advantage lies in its suitability for modular joint activation. Thus, if the active system can be applied to the commercial orthosis simply by adding or replacing a limited number of parts, the resulting versatility provides both an economic advantage and a practical advantage, as the orthosis can, for example, be adapted for rehabilitation alone. 4. Energy source A recurrent problem with active orthoses is the lack of a compact, lightweight energy source that can be installed on the device. Carrying the energy source is thus critical, and jeopardizes the structure's wearability. In view of the reliability and ease of maintenance of pneumatic actuation and the fact that compressed air sources are more commonly used than hydraulic ones and provide a better powerto-weight ratio than electrical actuation, it was decided to use a pneumatic source in the current solution. At this prototype stage of design, what kind of portable compressor could be connected to the exoskeleton has not been analysed and evaluated. Though such a solution is doubtless feasible, it is beyond the scope of the current investigation, which aims at reaching the desired gait performance for the orthosis, and would have shifted emphasis to optimising the bulk and dimensions of the generator and its supply source. This latter analysis can be carried out after the desired performance has been achieved, and the required pneumatic energy has been assessed and, consequently, minimized. For this ®rst prototype, performance was thus evaluated only with the system connected to a ®xed external supply source. 5. Design solutions The angles and corresponding torques of the joints to be moved must be evaluated before any decision can be made regarding type of actuation and kinematic layout. Trends for the torques applied to hip and knee joints during normal gait are shown in Fig. 3 [21]. As an initial rough approximation, these data can be used in sizing the system. A number of solutions were evaluated which involved the use of innovative components [23], as well as conventional pneumatic units such as rotary and linear actuators. The two solutions that were actually constructed and tested both use linear double-acting pneumatic cylinders, and dier in the type of layout employed to transmit the force generated by the piston to the joint concerned (Fig. 4).
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In the ®rst solution, the ends of the actuator are constrained via hinges to the top and bottom segments. In this way, the piston traction and thrust forces create the desired torques at cylindrical hinge level (Fig. 4a). An iterative procedure is necessary in order to achieve optimal use of the force that can be exerted by the actuator, bearing in mind both dimensional compatibility and the general requirements of minimum weight and bulk. The lever arm of this force must thus be selected on the basis of the best trade-o between a small lever arm, which ensures minimum bulk for the entire structure, and a large lever arm which makes it possible to exert more torque and limit actuator diameter. The idea for the layout shown in Fig. 4b springs from observation of the locomotor apparatus. The double-end-rod cylinder simulates the human muscle, while the chain and sprocket substitutes tendons and patella. Piston rod movement puts tension on the chain which, maintained taut and in line with the cylinder by the two chain tensioners, generates the required torque on the sprocket positively connected to the top segment. This solution is suitable for knee joint actuation, but cannot be used to move the hip joint because it would not adapt well to the femoral and trunk segments, while its excessive bulk would interfere with wearability.
Fig. 3. Movements applied to hip and knee joints.
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6. Logic and control system de®nition The management procedure was de®ned on the basis of the gait pattern described above. Considerable care was devoted to this stage, given that the system must be worn by a disabled subject who has no sensory perception in the lower limbs.
Fig. 4. Knee actuation system.
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The GEMMA (``Guide d'Etudes des Modes de Marches et d'ArreÃts'') method [24] was used to identify the various operating ®elds and to satisfy the requirements indicated earlier. Three dierent operating ®elds were identi®ed: normal starts and stops, emergency or failure stops and regulation. These groups are indicated in the management chart shown in Fig. 5. In the event of nonnormal operation, the general objective is to ensure that both limbs can always reach extension in order to guarantee a stable erect position. In this con®guration and under worst-case conditions, the system is a conventional passive orthosis which the patient is still able to manage. Normal starts and stops
Emergency and/or failure stops
normal operation takes place only if a start signal is present. The manager interprets the absence of this signal as a stop signal and the system comes to a halt at the end of the cycle, or in other words in the upright position.The maximum orthosis ¯exion angle is a function of the cylinder rod stroke, while the ¯exion and the extension speeds can be varied separately by means of ¯ow regulators.In addition, the system can be set up for either step-by-step or continuous cycle operation. As steps are rigidly cadenced in the latter case and do not depend on patient consent, it is clear that this type of management is more delicate and suitable only for trained patients who are able to coordinate the various movements without diculty. During the experiments, each step could be initiated only after the subject sent an enabling signal by means of a pushbutton. the most dangerous case is that in which a limb remains in ¯exion, as the subject could lose his/ her balance even if he/she still has at least three
Fig. 5. Management ¯ow - chart.
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Regulation
points of support. In this emergency condition, the system must be returned to the erect position. For this purpose, the status of the emergency signal is monitored at each step of the program. Should the emergency signal be present at any moment of the cycle, both limbs will be extended. A reset signal returns the system to the initial conditions, corresponding to the ®rst step. two regulation procedures are envisaged. The ®rst, which is entered directly in a main software routine, establishes the knee joint's maximum ¯exion angle. The second procedure makes it possible to synchronize the knee's ¯exion and extension speeds by means of the ¯ow regulators located in the discharge chamber of each cylinder. The orthosis must not be worn during the latter procedure, and for safety reasons the regulation is implemented in a dedicated routine which is separate from the main software program.
7. Electro-pneumatic layout The electro-pneumatic circuit is shown in Fig. 6. The type of solenoid valve used is dictated by safety considerations. Using normally open unistable valves (V2, and V3) to supply the lower cylinder chambers ensures that the knees will be extended if the electrical power supply fails. It also makes it possible to manage emergency events on two dierent levels: at the software level by sending a signal to the valves, and at the hardware level by cutting o electrical power supply. For knee ¯exion and for hip ¯exion and extension, normally closed valves (V1, V5, V6 and V4) were used. The ®gure shows the Reed units (R) which limit cylinder stroke and hence the ¯exion angle, and the ¯ow regulators (FR) used to control speed. The current need to test dierent control logics with maximum ¯exibility was satis®ed by using a programmable logic controller for cycle management. The system can be controlled directly by the user by means of buttons positioned on the crutches or by an operator using a push-button panel. At the preliminary design stage, which aimed at determining the feasibility of the active orthosis, movement was achieved without continuous speed control by means of pneumatic cylinders controlled by digital valves, as depicted in Fig. 6. This actuation does not make it possible to reproduce the physiological angle
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faithfully, as only an approximate pattern can be obtained. In addition, only the knee ¯ex/extension corresponding to the swing stage was considered, while the lower ¯ex/extension in the stance phase was neglected. Cycle control logic is managed by a FESTO FPC 101 PLC, and is described in the Grafcet shown in Fig. 7a, where two of the three pushbutton-selected operating modes considered in the ¯ow-chart in Fig. 5 are represented, viz., the normal gait cycle and regulation. The emergency management cycle is a higherlevel Grafcet, i.e., it can be activated at any stage of the cycle. For the normal gait cycle (Fig. 7c), there are two dierent operating cycles, which dier in how hip and knee movements are coordinated in time. In the ®rst case (cycle 1), hip activation takes place after knee ¯exion, with a delay set through timer T3. In cycle 2, on the other hand, the movements are simultaneous. The notation used for the transition signals and the command signals is the same
Fig. 6. Electro-pneumatic layout.
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as in the pneumatic diagram given in Fig. 6. In particular, valve actuation commands are denoted by lower case letters with the superscript +. Walking proceeds in step by step mode, with each lower limb moving forward only when the enabling signal is sent by means of pushbutton m. Finally, the regulation procedure entered directly in the main software routine makes it possible to set and check the knee joint's maximum ¯exion angles by moving the limb manually.
Fig. 7. Grafcet of the management program.
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8. Experimental tests Experimental testing followed system development closely, as design could not be carried out without step-by-step performance checks. Experimental tests included: . Tests with no user . Tests with healthy subjects . Tests with paraplegic subjects The ®rst two sets of test are a necessary prelude to the third, which is obviously the most interesting and the most meaningful. For the structure in question, it is clear that the experimental tests are not simply a means of collecting numerical data, but are also an occasion for gathering information about the orthosis wearer's perceptions and suggestions. While graphs provide clear-cut and immediate comparison with published results, perceptions, though subjective, can provide an idea of how the structure actually works and how it could be improved through changes in design variables. During tests, rotations and applied torques versus cylinder pressure were determined for each joint. In addition, all tests were ®lmed in order to correlate numerical data and images and thus identify the parameters which are indicative of gait quality. Kinematic magnitudes and exchanged forces were measured using appropriate transducers. Transducers were chosen so as to be as little invasive as possible, both to maintain wearability and because of the particular structure of the orthosis. To determine hip and knee joint rotations, rotary potentiometers were installed directly on the hinge axis of rotation. Cylinder chamber pressures were measured in order to evaluate applied drive torque, while Reed stroke limiter signals and the step initiation enabling signal were also recorded. The ®rst test addressed gait symmetry and repetitiveness, which, slight dierences apart, were found to be good. Though the fact that movement is periodic makes it possible to limit analysis of the magnitudes involved to a period equal to that of each step, the test duration must be longer to enable the user to establish regular motion and evaluate his perceptions. Tests were conducted with dierent values for supply pressure and ¯ex/ extension angles and speeds. Parameters were varied individually in accordance with information provided by the user and the gait achieved.
8.1. Tests with no user The ®rst tests were carried out with no user, i.e. the orthosis was not worn. Stability and limb forward movement, which are normally provided by the patient, were in this case obtained by means of an external support structure and
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a rotary motor applied at the hip joint (Fig. 8). Given the type of test involved, the system control circuit used was simpler than that described above. Using the structure as a bipedal robot made it possible to verify its feasibility, and to evaluate and modify the various gait cycle parameters. In addition, it revealed the limitations of the design solution shown in Fig. 4a as compared to that in Fig. 4b. In the ®rst con®guration, the actuator position varies according to
Fig. 8. Experimental test with no user.
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knee ¯exion angle, as does applied torque. Most importantly, however, this movement is critical for patient safety, and overall dimensions increase along with knee ¯exion angle. By contrast, the solution shown in Fig. 4b is more compact and, once engineered, is also better in terms of cosmesis because it does not prevent the orthosis from being worn under the patient's clothing. This solution was thus used to activate the knee hinge joint, while the solution in Fig. 4a was used for the hip joint, where rotational angles are smaller. 8.2. Tests with healthy subjects These experimental tests were conducted with the orthosis worn by a healthy subject who simulated lack of muscle control to the greatest possible extent. These test conditions can be considered sucient during a preliminary evaluation of the system carried out in order to assess its reliability and establish the gait pattern that can be achieved, but they entail limitations which must be borne in mind if we are to avoid glaring errors in evaluating and interpreting results. Nevertheless, this intermediate step between operating the structure as a bipedal robot and ®tting it to a paraplegic subject was considered to be a useful means of gauging the structure's potential and reliability, and of gaining familiarity with it. Indeed, the fact that a healthy subject has muscle control and skin sensitivity in the lower limbs makes it possible to test the orthosis under dierent conditions using various gait parameters. These tests also addressed the feasibility of not activating hip joint rotation and thus leaving forward limb movement to the user, as with passive orthoses. The presence of the reciprocating cable and the fact that the knee can ¯ex during the swing phase reduces the eort required of the patient and makes this solution interesting since the corresponding actuators could be eliminated. As regards pressures, it was found that the pressure in the lower chamber, which determines knee extension, never drops to zero. This behaviour is linked both to the necessary speed regulation and to the fact that the cylinder does not make use of the entire rod stroke during gait. This promotes safety, as the system is better prepared to guarantee a stable upright position in the event of an emergency. By contrast, pressure in the upper chamber rises abruptly as soon as the enabling signal is received, and stays close to the supply pressure throughout the ¯exion phase. Figs. 9 and 10 show the angles of the two joints of the left limb versus time, while Fig. 11 shows the torque acting on the knee, which is proportional to the dierence between the pressures in the cylinder's upper and lower chambers. In all the three graphs, the enabling signal sent by the patient in order to initiate the step is represented by a dashed line. For the knee joint, a slight unforeseen ¯exion occurs during the stance phase, which in Fig. 9 corresponds roughly to the time periods 1±6.5 s and 9.5±15.5 s. This ¯exion takes place at the time the right limb is ¯exed, and is due to the type of pneumatic circuit used. It was not eliminated because it is in any case close to the physiological behaviour (Fig. 1).
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Fig. 9. Left knee angle.
In the graph shown in Fig. 10, which refers to the hip joint, the positive angles denote ¯exion while the negative angles denote extension. It is clear from the graph that hip and knee ¯exion are simultaneous; this shows that the operations preceding the swing phase are well coordinated, and minimizes energy expenditure. The torque curve in Fig. 11 has a number of signi®cant characteristics. First, it should be emphasized that this torque is equal to the drive torque acting on the joint only when ¯exion is greater than zero. During full extension, on the other hand, the actuator piston is at the end of stroke against the end plate, and part of the thrust force is unloaded against this constraint. The torque assumed in this situation, approximately ÿ60 Nm at 2 s, is maintained above the physiological value in order to prevent system collapse and provide a sucient margin of safety. Obviously, this torque can be varied as a function of supply pressure and in relation to the size of the pneumatic actuator used. When the right limb is ¯exed at 2.5 s, there is a reduction in absolute torque
Fig. 10. Left hip.
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corresponding to a slight ¯exion in Fig. 9, even though none of the left actuator valves are activated. The torque value is in any case sucient to guarantee stability, rising to the maximum value at approximately 6.5 s. Following the ¯exion command (at approximately 6.5 s), the absolute torque value drops, maintaining a modest positive value during the ¯exion movement. Similarly, it maintains a small negative value starting from the extension command at approximately 8 s, rising to the absolute maximum once full extension is reached. Using a force above the physiological value is generally dangerous, since in a paralyzed subject with no antagonist muscle force it could cause limb hyperextension with consequent ligament injury. In the case in question, this possibility does not in fact exist, as the cylinder is at the end of stroke at the time of maximum negative torque. A mechanical stop is thus provided which prevents 08 from being exceeded in extension. Subsequent tests indicated that even a small variation in the restrictors` ¯ow coecient resulted in dierent ¯ex/extension speeds, and hence in qualitatively dierent gaits. It was also observed that the absolute value of the extension moment drops when supply pressure is reduced, while the changes in the ¯exion torque are negligible. This is positive, as it is thus possible to come closer to the maxima and minima of the corresponding physiological values, and in particular to reduce the extension moment without causing appreciable changes in the ¯exion moment, which, as in the previous tests, remains within normal gait limits. 8.3. Tests with a paraplegic subject In experimental tests performed with a paraplegic subject having a T3 lesion level, the solution without hip actuation was found to be too dicult because of the increased weight of the knee actuators. By contrast, good results were achieved with the ``hybrid'' solution, where part of the energy is provided by the
Fig. 11. Torque trends on the knee joint.
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user and part by the pneumatic actuators. This solution calls for good coordination between the user and the phases of the gait cycle, which can be developed through training. An image sequence ®lmed during a test is shown in Fig. 12, while curves for the recorded magnitudes are given in the following ®gures. Tests were carried out with crutches, a walker and the parallel bars to maintain stability. As the parallel bars provide a greater sense of security, they were used in the majority of tests. It will be noted that the images in Fig. 12 coincide with the stick diagrams in Fig. 2 to a considerable extent. The cycle starts with left leg swing at approximately t = 45 s and ends at about t = 52 s. In particular, Fig. 12a (t = 45 s) shows the pre-swing of the left leg: this is still a double support phase, but body weight is almost supported by the stance leg, i.e., the right one. In Fig. 12b (t = 46 s), the swing leg passes the stance foot and the knee is still ¯exed, while the left hip angle is about zero. At t = 47 s (Fig. 12c), the left knee is completely extended, but the swing motion has not yet been accomplished, in fact the left hip angle is 108. The maximum stride length, corresponding to maximum hip ¯exion, is achieved at t = 48 s when the left foot touches the ground. Preparation for the swing phase of the right limb is depicted in Fig. 12e: the left arm is shifted ahead so that the body's center of gravity is also moved forward. This movement also causes the left foot tip to lower and the right heel to lift. Photographs of Fig. 12f±h, corresponding to t = 50, 51 and 52 s, respectively, are related to the swing phase of the right leg. The comparison between this experimental test and the one with the healthy subject shows a good level of correspondence between the dierent ¯exion angle curves. The slight dierences can be attributed to a number of factors, of which the most important is the fact that the orthosis was worn by a dierent user. For instance, the knee stance phase ¯exion (Fig. 13) which occurred in the earlier tests has now disappeared, and as a consequence the knee joint torque curve also diers (Fig. 14). This is due to the fact that the weight of the paraplegic patient is considerably lower than that of healthy user, and in this test the weight unloaded through the arms is probably higher. The dierence between left and right hip rotation range (Fig. 15), and hence stride length, is due to the user, whose left latissimus dorsi muscle injuries prevent him from exercising the same trunk extension and thus from making symmetrical strides. The maximum hip ¯exion and extension angles occur only when knee ¯exoextension is ®nished and the user extends the trunk and loads the rear cable. The values of the torque supplied by the pneumatic actuators are about 50% of the physiological value, as depicted in Fig. 16. In the test described here, activation of the hip and knee actuators is simultaneous. Dierent sequences of the valve command signals were used, but this pattern permitted a good gait and was also considered by the wearer to be the best.
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Fig. 12. Sequential images of a gait cycle.
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Fig. 13. Knee rotation angles.
Fig. 14. Hip rotation angles.
Fig. 15. Right knee torque.
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Fig. 16. Right hip torque.
9. Conclusions The orthosis presented herein is intended to assist locomotion in paraplegic subjects, including those with high-level lesions. In addition to being suitable for rehabilitation, it can be used on a daily basis in structured environments such as the home or workplace. The device is ideal for the rehabilitation stage, given the structure's modularity and the extremely ¯exible means used to regulate gait characteristics. Future work will address improvements in wearability and the gait cycle. As regards wearability, we intend to reduce the structure's weight and bulk, re-assess actuator placement and, if necessary, investigate other types of actuator. Changes in the gait cycle will focus on reducing physiological energy expenditure and will include the introduction of an ankle joint or a device capable of lifting the user during the one-legged stance phase so that the swinging limb will not strike the ground. Finally, a gait cycle which resembles the normal cycle in all respects, at least in the sagittal plane, can be achieved by using proportional interfaces for force control and for coordinating movements and the torques exerted by the actuators. Acknowledgements The research was supported by the Istituto Superiore della SanitaÁ. The authors would like to thank Steeper Europe BVBA for providing the ARGO. References [1] Nene AV, Hermens HJ, Zilvold G. Paraplegic locomotion: a review. Spinal Cord 1996;34:507±24.
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[2] Jasper J, Peeraer L, Van Petegem W, Van der Perre G. The use of an advanced reciprocating gait orthosis by paraplegic individuals: a follow-up study. Spinal Cord 1997;35:289±585. [3] Sykes L, Edwards J, Powell ES, Ross ER. The reciprocating gait orthosis: long-term usage patterns. Arch Phys Med Rehabil 1995;76:779±83. [4] Ogilvie C, Bower P, Rowley DI. The of paraplegic physiological bene®ts rthotically aided walking. Paraplegia 1993;31:111±5. [5] Stallard J, Major RE, Patrick JH. A review of the fundamental design problems of providing ambulation for paraplegic patients. Paraplegia 1989;27:70±5. [6] Lotta S, Fiocchi A, Giovannini R, Silvestrin R, Tesio L, Raschi A, Macchia L, Chiapatti V, Zambelli M, Tosi C, Baratta S, Franceschini M. Restoration of gait with orthosis in thoracic paraplegia: a multicentric investigation. Paraplegia 1994;32:608±15. [7] Glaser RM, Sawka MN, Wilde SW, Woodrow BK, Suryprasad AG. Energy cost and cardiopulmonary responses for wheelchair locomotion and walking on the tile and on carpet. Paraplegia 1981;19:220±6. [8] Massucci M, Brunetti G, Betti L, Franceschini M. Walking with the advanced reciprocating gait orthosis (ARGO) in thoracic paraplegic patients: energy expenditure and cardiorespiratory performances. Spinal Cord 1998;36:223±7. [9] Bernardi M, Canale I, Castellano V, Di Filippo L, Felici F, Marchetti M. The eciency of walking paraplegic patients using a reciprocating gait orthosis. Paraplegia 1995;33:409±15. [10] Solomov M, Shoji H, D'Ambrosia R, Douglas R. Electromechanical walking system for paraplegics. In: Proceedings of the IEEE 7th Annual Conference of the Engineering in Medicine and Biology Society. 1995. p. 4±7. [11] Vukobratovic M, Hristic D, Stojiljkovic Z. Development of active anthropomorphic exoscheletons. Medical and Biological Engineering 1974;12:66±80. [12] Rabischong P. Robotics for the Handicapped. In: Proceedings IFAC Control Aspects of Prosthetics and Orthotics (Columbus, OH, USA): May 7±9. 1982. p. 163±7. [13] Rabischong E, Sgarbi F, Rabischong P, Detriche JM, Pinguet N, Riwan A. Control and command of a six degree of freedom active electrical orthosis for paraplegic patient. In: Proceedings IEEE Inter. Workshop on Intelligent Robots and Systems, IROS '90 (Tsuchiura, Japan), July 3±6. 1990. p. 987±91. [14] Miyamoto H, Israel I, Miyamoto H, Mori S, Sand A, Sakurai Y. Approach to a powered orthosis for paralysed lower limbs. In: Proceedings ICAR '85. 1985. p. 458±551. [15] Miyamoto H. Powered orthosis for lower limbs Ð its structure and control. J Robotics and Mechatronics 1993;5(5):443±7. [16] Saunders JB, Inman VT, Eberhart HD. The major determinant in normal and pathological gait. The J of Bone and Joint Surgery 1953;35-A(3):543±58. [17] Beckett R, Chang K. An evaluation of the kinematics of gait by minimum energy. J Biomechanics 1968;1:147±59. [18] Belforte G, Sorli M, Gastaldi L. Applicazioni di dispositivi pneumatici al comando di tutori attivi. In: Proceedings Trasmissioni di Potenza (Assago, Italy): June 10±11. 1997. p. 118±31. [19] Belforte G, Sorli M, Gastaldi L. Active orthosis for rehabilitation and passive exercise. In: Proceedings Fourth International Conference on Simulation in Biomedicine, BIOMED '97 (Acquasparta, Italy): June 9±11. 1997. p. 199±208. [20] Belforte G, Gastaldi L, Sorli M, Manuello Bertetto A, Mazzeo M. Active orthosis: experiments with a healthy subject. In: Proceedings 8th International Workshop on Robotics, RAAD '99 (Munchen, Germany): June 17±19. 1999. p. 317±22. [21] Vaughan CL, Davis BL, O'Connor JC. Gait Lab. Champain, IL: Human Kinetics Publishers. [22] Steeper Advanced Reciprocating Gait Orthosis. Fabrication and Fitting Manual. Steeper. [23] Ferraresi C, Franco W, Manuello Bertetto A. Straight ®bbers pneumatic muscle: an actuator with high traction force. In: Proceedings 6th Scandinavian International Conference on Fluid Power, SICFP '99. (Tampere, Finland): May 26±28. 1999. p. 787±98. [24 Bouteille D, Belforte G. Automazione ¯essibile elettropneumatica e pneumatica. Milano, Italy: Tecniche Nuove, 1987.
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Glossary Anterior: Toward the front of the body (i.e., the toes are anterior to the heel) Extension: straightening movement of a joint, whereby the two contiguous segments form a more or less obtuse angle (for the hip joint, for example, extension brings the thigh posterior to the trunk) Flexion: bending movement of a joint, whereby the two contiguous segments form a more or less acute angle (for the hip joint, for example, ¯exion brings the thigh anterior to the trunk) Dorsi¯exion: lifting the tip of the foot or hand Plantar ¯exion: lowering the tip of the foot Inferior: with the body erect, the position toward the feet Hyperextension: extension beyond the anatomical position Lateral: farther from the medial plane Medial: toward the midline of the body, or on the median plane Frontal plane: plane dividing body into an anterior portion and a posterior portion. Also called coronal plane Sagittal plane: plane dividing the body into a right and left half Transverse plane: plane dividing the body into a superior portion and an inferior portion Posterior: towards the back of the body (i.e., the heel is posterior to the toes) Rotation: movement whereby a bone or a series of bones rotates around an axis passing through a joint Superior: with the body erect, toward the head Stance phase: gait cycle phase when the foot is on the ground Swing phase: gait cycle phase when the foot is lifted o the ground
Pneumatic active gait orthosis G. Belforte, L. Gastaldi*, M. Sorli Department of Mechanics, Politecnico di Torino, c.so Duca degli Abruzzi, 24 10129 Torino, Italy Received 30 July 1999; received in revised form 8 February 2000; accepted 1 March 2000
Abstract The paper describes design, construction and experimental testing of an active gait orthosis intended to assist locomotion in paraplegic subjects. Design speci®cations were formulated on the basis of an analysis of the context in which the device will be used. Dierent joint activation solutions were then proposed and examined. Considerable care was devoted to de®ning the logic and control system and to implementing the electropneumatic circuit, given that the system must be worn by a disabled subject who has no sensory perception in the lower limbs. Experimental tests, which made it possible to determine system performance at each step of development, were carried out with no user, then with a healthy user and ®nally with a paraplegic user. Experimental results consisting of graphs and photographic images are presented and discussed. 7 2000 Elsevier Science Ltd. All rights reserved. Keywords: Pneumatic; Active orthosis; Paraplegia
1. Introduction Signi®cant resources continue to be invested in the development of orthoses for functional gait restoration in paraplegic patients [1]. There can be no doubt that being able to maintain an upright position is psychologically important. Having to use a wheelchair in surroundings which are normally designed for people who can walk, not only makes everyday life immeasurably more complicated, but also * Corresponding author. Tel.: +39-011-564-6939; fax: +39-011-564-6999. E-mail address: [email protected] (L. Gastaldi). 0957-4158/01/$ - see front matter 7 2000 Elsevier Science Ltd. All rights reserved. PII: S 0 9 5 7 - 4 1 5 8 ( 0 0 ) 0 0 0 1 7 - 9
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aects personal relationships and communication [2,3]. From the physiological standpoint, a number of therapists maintain Ð though without experimental documentation Ð that the upright position promotes better body system function, prevents bedsores, and combats osteoporosis, while in Ref. [4], the physiological bene®ts of an upright position, at least in the short term, are claimed to be not so great. In approaching this problem, however, it must be borne in mind that restoring locomotion is not generally seen by the disabled as a fundamental need, and in any case depends on how it is accomplished. After the initial enthusiasm at being able to stand and walk again, in fact, the patient is faced with serious psychological and physiological problems, involving the cardio-vascular, the urological and the respiratory apparatus, in comparison with which walking becomes a secondary consideration. In addition, the patient must cope with the diculties associated with locomotion itself and with the orthopaedic aids, including high energy expenditure, diculty in putting on the orthosis, poor cosmesis, and unnatural gait [5]. All this leads to a feeling of disappointment, which in most cases at least [3,6], causes the patient to abandon the orthosis after initial rehabilitation at the orthopaedic clinic. Thus, the functional need to make the upright position part of the disabled subject's daily life once again is oset by an insucient patient motivation, which is partly justi®ed by the fact that the wheelchair is an excellent means of covering considerable distances with a modest energy expenditure comparable to that of normal gait [7]. Currently, the only means whereby the upright position and locomotion can be recovered are passive orthoses. These devices make up for the inability of paralysed postural muscles to contract by stiening the lower limbs. The passive orthoses which provide best performance are the reciprocating types, which enable the user to assume an upright posture without stabilising aids and can also be used by patients with high-level lesions. Walking, on the other hand, calls for the use of crutches or a walking frame and, as it takes place with the legs fully extended, is extremely tiring [8,9]. Other prototype systems involve the use of functional electrical stimulation (FES), but the large number of channels and the resulting muscle fatigue limit the distance that can be covered. Hybrid systems requiring fewer channels are more promising, as FES is associated with a support exoskeleton [10]. Another approach is that of using active orthoses, i.e., orthoses in which the energy required for locomotion is provided by an external energy source. A prototype system of this kind was constructed by Vukobratovic [11]; here, system stabilization and the fact that all lower limb joints are activated in the sagittal plane makes this structure heavy, bulky and complex from the control standpoint. A more recent investigation of walking machines was conducted by Rabischong et al. [12,13] who proposed a master slave control system for training paraplegic patients to walk again. This system consists of a passive orthosis worn by the physiotherapist and an active orthosis worn by the patient, which is equipped with actuators so that the trainer's movements are faithfully reproduced. A third active
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orthosis was proposed in Refs. [14,15] by Miyamoto et al. Hip and knee joints are activated hydraulically and equilibrium control is maintained by the patient. Generally, these solutions cannot yet be regarded as viable alternatives to the wheelchair. They are used for therapeutic purposes in order to achieve long-term bene®ts, or for mobility in structured environments such as the home, workplace or school. Higher walking speeds and lower eort levels are two requisites which would clearly encourage more extensive orthosis use, both by those who already wear an orthosis on a regular basis, and by those who do not use their orthosis because of the excessive energy expenditure involved. The active orthosis developed at the Department of Mechanics of the Politecnico di Torino was designed to provide these requisites and, where possible, those indicated in Ref. [5] as essential, viz., cosmesis, wearability, relative simplicity, etc. As these parameters are fundamental to the structure's acceptance by paraplegic patients, the trade-o which best satis®es the user must be identi®ed. Formulation of design speci®cations was preceded by an analysis of the operation of passive orthoses (Hip Guidance Orthosis, Reciprocating Gait Orthosis, Advanced Reciprocating Gait Orthosis) and active prototypes, while consideration was also given to minimizing the energy cost of locomotion [16,17]. Throughout orthosis development, moreover, rehabilitation therapists and paraplegic patients provided valuable feedback. The prototype consists of a passive support structure and a pneumatic actuation system which powers and controls the hip and knee joints. Possible design solutions were discussed in Refs. [18,19], where initial tests carried out on the orthosis without a wearer were also presented. The results of tests conducted with healthy subjects were presented in Ref. [20]. This paper presents the ®nal version of the active orthosis prototype, the actuation and control system, and the results of experimental tests performed with a paraplegic user. 2. Gait pattern The desirable properties of an orthosis can be summed up as providing the ability to walk, with energy savings, the ability to provide near-normal gait and a highly wearable structure. It is important to evaluate which degrees of freedom must be introduced and which must be blocked in order to have satisfactory gait with the minimum number of actuators. To do so, we assessed the in¯uence of each main characteristic of reciprocating gait [16] on the trajectory of the body's center of gravity. This analysis was carried out only on the sagittal plane, as it is there that the larger movements occur. The hip, knee and ankle rotation angles in the sagittal plane resulting from analysis of physiological gait are shown in Fig. 1 [21]. Hip ¯exion permits the lower limbs to move forward in alternation; at heel
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strike, the hip is in ¯exion and the trunk is situated to the rear of the foot that has been swung forward, while in the subsequent stance phase the hip is fully extended. Knee ¯exion occurs both during the stance phase, where it minimizes the center of gravity's vertical excursion, and during the swing phase where it prevents the foot from hitting the ground. Thus, knee ¯exion essentially reduces energy consumption. Plantar ¯exion takes place at various times during gait: when the heel contacts the ground and the ankle moves downwards and to the rear, at the terminal swing phase to prevent the foot from dragging, and at the end of the stance phase when the body's entire weight must be propelled forward. From these considerations, ¯anked by an assessment of costs versus bene®ts, it follows that the hip and knee joints in the sagittal plane are indispensable for near-normal gait and for reducing the excursion of the body's center of gravity and, consequently, energy expenditure.
Fig. 1. Hip, knee and ankle joint rotations in the sagittal plane.
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Stick diagrams for evaluating the gait pattern that can be achieved with the assumptions made above are shown in Fig. 2 for the various phases, both in the sagittal and transversal plane. The balance is maintained by the user with a pair of crutches. Diagram A in Fig. 2 shows the instant of initial contact between the right foot and the ground, which corresponds to the beginning of the stance phase and of the double support period (the hatched footprints indicate contact with the ground). The subsequent loading response phase (B) prepares for the left limb's swing motion: body weight is transferred from the left to the right limb, and the left crutch is positioned anteriorly in order to ensure balance in the subsequent phases. In the midstance phase (C), the swinging limb moves ahead of the stance limb. Forward movement is produced by ¯exion of the left hip, and by ¯exion of the knee, which prevents the foot from hitting the ground. Diagram D shows the terminal moment of the single limb stance phase, when the left knee is extended prior to contact with the ground. Diagram E showing preparation for the swing phase and diagrams F, G and H, which correspond respectively to the initial swing, midswing and terminal swing phases, are symmetrical to the previous diagrams, given that they represent the right leg's swing motion. 3. Exoskeleton The prototype's support structure is a commercial passive reciprocating gait orthosis, modi®ed to introduce new degrees of freedom and accommodate the actuators. As only two-dimensional gait was assumed, the knee and hip joints consist of cylindrical hinges and ankle rotation is prevented by immobilizing the joint with an Ankle Foot Orthosis (AFO). The choice of a commercial structure such as the ARGO (Advanced Reciprocating Gait Orthosis) [22] made it possible to use a structure which has already been optimized, particularly as regards the hip hinge joints and the rear
Fig. 2. Gait phases with the new active orthosis.
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cable. The latter rigidly connects the two hips to generate a reciprocating movement: in other words, extension of one hip empowers ¯exion of the contralateral hip and vice versa. As regards psychological acceptance, most paraplegics are already familiar with the structure. A further advantage lies in its suitability for modular joint activation. Thus, if the active system can be applied to the commercial orthosis simply by adding or replacing a limited number of parts, the resulting versatility provides both an economic advantage and a practical advantage, as the orthosis can, for example, be adapted for rehabilitation alone. 4. Energy source A recurrent problem with active orthoses is the lack of a compact, lightweight energy source that can be installed on the device. Carrying the energy source is thus critical, and jeopardizes the structure's wearability. In view of the reliability and ease of maintenance of pneumatic actuation and the fact that compressed air sources are more commonly used than hydraulic ones and provide a better powerto-weight ratio than electrical actuation, it was decided to use a pneumatic source in the current solution. At this prototype stage of design, what kind of portable compressor could be connected to the exoskeleton has not been analysed and evaluated. Though such a solution is doubtless feasible, it is beyond the scope of the current investigation, which aims at reaching the desired gait performance for the orthosis, and would have shifted emphasis to optimising the bulk and dimensions of the generator and its supply source. This latter analysis can be carried out after the desired performance has been achieved, and the required pneumatic energy has been assessed and, consequently, minimized. For this ®rst prototype, performance was thus evaluated only with the system connected to a ®xed external supply source. 5. Design solutions The angles and corresponding torques of the joints to be moved must be evaluated before any decision can be made regarding type of actuation and kinematic layout. Trends for the torques applied to hip and knee joints during normal gait are shown in Fig. 3 [21]. As an initial rough approximation, these data can be used in sizing the system. A number of solutions were evaluated which involved the use of innovative components [23], as well as conventional pneumatic units such as rotary and linear actuators. The two solutions that were actually constructed and tested both use linear double-acting pneumatic cylinders, and dier in the type of layout employed to transmit the force generated by the piston to the joint concerned (Fig. 4).
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In the ®rst solution, the ends of the actuator are constrained via hinges to the top and bottom segments. In this way, the piston traction and thrust forces create the desired torques at cylindrical hinge level (Fig. 4a). An iterative procedure is necessary in order to achieve optimal use of the force that can be exerted by the actuator, bearing in mind both dimensional compatibility and the general requirements of minimum weight and bulk. The lever arm of this force must thus be selected on the basis of the best trade-o between a small lever arm, which ensures minimum bulk for the entire structure, and a large lever arm which makes it possible to exert more torque and limit actuator diameter. The idea for the layout shown in Fig. 4b springs from observation of the locomotor apparatus. The double-end-rod cylinder simulates the human muscle, while the chain and sprocket substitutes tendons and patella. Piston rod movement puts tension on the chain which, maintained taut and in line with the cylinder by the two chain tensioners, generates the required torque on the sprocket positively connected to the top segment. This solution is suitable for knee joint actuation, but cannot be used to move the hip joint because it would not adapt well to the femoral and trunk segments, while its excessive bulk would interfere with wearability.
Fig. 3. Movements applied to hip and knee joints.
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6. Logic and control system de®nition The management procedure was de®ned on the basis of the gait pattern described above. Considerable care was devoted to this stage, given that the system must be worn by a disabled subject who has no sensory perception in the lower limbs.
Fig. 4. Knee actuation system.
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The GEMMA (``Guide d'Etudes des Modes de Marches et d'ArreÃts'') method [24] was used to identify the various operating ®elds and to satisfy the requirements indicated earlier. Three dierent operating ®elds were identi®ed: normal starts and stops, emergency or failure stops and regulation. These groups are indicated in the management chart shown in Fig. 5. In the event of nonnormal operation, the general objective is to ensure that both limbs can always reach extension in order to guarantee a stable erect position. In this con®guration and under worst-case conditions, the system is a conventional passive orthosis which the patient is still able to manage. Normal starts and stops
Emergency and/or failure stops
normal operation takes place only if a start signal is present. The manager interprets the absence of this signal as a stop signal and the system comes to a halt at the end of the cycle, or in other words in the upright position.The maximum orthosis ¯exion angle is a function of the cylinder rod stroke, while the ¯exion and the extension speeds can be varied separately by means of ¯ow regulators.In addition, the system can be set up for either step-by-step or continuous cycle operation. As steps are rigidly cadenced in the latter case and do not depend on patient consent, it is clear that this type of management is more delicate and suitable only for trained patients who are able to coordinate the various movements without diculty. During the experiments, each step could be initiated only after the subject sent an enabling signal by means of a pushbutton. the most dangerous case is that in which a limb remains in ¯exion, as the subject could lose his/ her balance even if he/she still has at least three
Fig. 5. Management ¯ow - chart.
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Regulation
points of support. In this emergency condition, the system must be returned to the erect position. For this purpose, the status of the emergency signal is monitored at each step of the program. Should the emergency signal be present at any moment of the cycle, both limbs will be extended. A reset signal returns the system to the initial conditions, corresponding to the ®rst step. two regulation procedures are envisaged. The ®rst, which is entered directly in a main software routine, establishes the knee joint's maximum ¯exion angle. The second procedure makes it possible to synchronize the knee's ¯exion and extension speeds by means of the ¯ow regulators located in the discharge chamber of each cylinder. The orthosis must not be worn during the latter procedure, and for safety reasons the regulation is implemented in a dedicated routine which is separate from the main software program.
7. Electro-pneumatic layout The electro-pneumatic circuit is shown in Fig. 6. The type of solenoid valve used is dictated by safety considerations. Using normally open unistable valves (V2, and V3) to supply the lower cylinder chambers ensures that the knees will be extended if the electrical power supply fails. It also makes it possible to manage emergency events on two dierent levels: at the software level by sending a signal to the valves, and at the hardware level by cutting o electrical power supply. For knee ¯exion and for hip ¯exion and extension, normally closed valves (V1, V5, V6 and V4) were used. The ®gure shows the Reed units (R) which limit cylinder stroke and hence the ¯exion angle, and the ¯ow regulators (FR) used to control speed. The current need to test dierent control logics with maximum ¯exibility was satis®ed by using a programmable logic controller for cycle management. The system can be controlled directly by the user by means of buttons positioned on the crutches or by an operator using a push-button panel. At the preliminary design stage, which aimed at determining the feasibility of the active orthosis, movement was achieved without continuous speed control by means of pneumatic cylinders controlled by digital valves, as depicted in Fig. 6. This actuation does not make it possible to reproduce the physiological angle
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faithfully, as only an approximate pattern can be obtained. In addition, only the knee ¯ex/extension corresponding to the swing stage was considered, while the lower ¯ex/extension in the stance phase was neglected. Cycle control logic is managed by a FESTO FPC 101 PLC, and is described in the Grafcet shown in Fig. 7a, where two of the three pushbutton-selected operating modes considered in the ¯ow-chart in Fig. 5 are represented, viz., the normal gait cycle and regulation. The emergency management cycle is a higherlevel Grafcet, i.e., it can be activated at any stage of the cycle. For the normal gait cycle (Fig. 7c), there are two dierent operating cycles, which dier in how hip and knee movements are coordinated in time. In the ®rst case (cycle 1), hip activation takes place after knee ¯exion, with a delay set through timer T3. In cycle 2, on the other hand, the movements are simultaneous. The notation used for the transition signals and the command signals is the same
Fig. 6. Electro-pneumatic layout.
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as in the pneumatic diagram given in Fig. 6. In particular, valve actuation commands are denoted by lower case letters with the superscript +. Walking proceeds in step by step mode, with each lower limb moving forward only when the enabling signal is sent by means of pushbutton m. Finally, the regulation procedure entered directly in the main software routine makes it possible to set and check the knee joint's maximum ¯exion angles by moving the limb manually.
Fig. 7. Grafcet of the management program.
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8. Experimental tests Experimental testing followed system development closely, as design could not be carried out without step-by-step performance checks. Experimental tests included: . Tests with no user . Tests with healthy subjects . Tests with paraplegic subjects The ®rst two sets of test are a necessary prelude to the third, which is obviously the most interesting and the most meaningful. For the structure in question, it is clear that the experimental tests are not simply a means of collecting numerical data, but are also an occasion for gathering information about the orthosis wearer's perceptions and suggestions. While graphs provide clear-cut and immediate comparison with published results, perceptions, though subjective, can provide an idea of how the structure actually works and how it could be improved through changes in design variables. During tests, rotations and applied torques versus cylinder pressure were determined for each joint. In addition, all tests were ®lmed in order to correlate numerical data and images and thus identify the parameters which are indicative of gait quality. Kinematic magnitudes and exchanged forces were measured using appropriate transducers. Transducers were chosen so as to be as little invasive as possible, both to maintain wearability and because of the particular structure of the orthosis. To determine hip and knee joint rotations, rotary potentiometers were installed directly on the hinge axis of rotation. Cylinder chamber pressures were measured in order to evaluate applied drive torque, while Reed stroke limiter signals and the step initiation enabling signal were also recorded. The ®rst test addressed gait symmetry and repetitiveness, which, slight dierences apart, were found to be good. Though the fact that movement is periodic makes it possible to limit analysis of the magnitudes involved to a period equal to that of each step, the test duration must be longer to enable the user to establish regular motion and evaluate his perceptions. Tests were conducted with dierent values for supply pressure and ¯ex/ extension angles and speeds. Parameters were varied individually in accordance with information provided by the user and the gait achieved.
8.1. Tests with no user The ®rst tests were carried out with no user, i.e. the orthosis was not worn. Stability and limb forward movement, which are normally provided by the patient, were in this case obtained by means of an external support structure and
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a rotary motor applied at the hip joint (Fig. 8). Given the type of test involved, the system control circuit used was simpler than that described above. Using the structure as a bipedal robot made it possible to verify its feasibility, and to evaluate and modify the various gait cycle parameters. In addition, it revealed the limitations of the design solution shown in Fig. 4a as compared to that in Fig. 4b. In the ®rst con®guration, the actuator position varies according to
Fig. 8. Experimental test with no user.
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knee ¯exion angle, as does applied torque. Most importantly, however, this movement is critical for patient safety, and overall dimensions increase along with knee ¯exion angle. By contrast, the solution shown in Fig. 4b is more compact and, once engineered, is also better in terms of cosmesis because it does not prevent the orthosis from being worn under the patient's clothing. This solution was thus used to activate the knee hinge joint, while the solution in Fig. 4a was used for the hip joint, where rotational angles are smaller. 8.2. Tests with healthy subjects These experimental tests were conducted with the orthosis worn by a healthy subject who simulated lack of muscle control to the greatest possible extent. These test conditions can be considered sucient during a preliminary evaluation of the system carried out in order to assess its reliability and establish the gait pattern that can be achieved, but they entail limitations which must be borne in mind if we are to avoid glaring errors in evaluating and interpreting results. Nevertheless, this intermediate step between operating the structure as a bipedal robot and ®tting it to a paraplegic subject was considered to be a useful means of gauging the structure's potential and reliability, and of gaining familiarity with it. Indeed, the fact that a healthy subject has muscle control and skin sensitivity in the lower limbs makes it possible to test the orthosis under dierent conditions using various gait parameters. These tests also addressed the feasibility of not activating hip joint rotation and thus leaving forward limb movement to the user, as with passive orthoses. The presence of the reciprocating cable and the fact that the knee can ¯ex during the swing phase reduces the eort required of the patient and makes this solution interesting since the corresponding actuators could be eliminated. As regards pressures, it was found that the pressure in the lower chamber, which determines knee extension, never drops to zero. This behaviour is linked both to the necessary speed regulation and to the fact that the cylinder does not make use of the entire rod stroke during gait. This promotes safety, as the system is better prepared to guarantee a stable upright position in the event of an emergency. By contrast, pressure in the upper chamber rises abruptly as soon as the enabling signal is received, and stays close to the supply pressure throughout the ¯exion phase. Figs. 9 and 10 show the angles of the two joints of the left limb versus time, while Fig. 11 shows the torque acting on the knee, which is proportional to the dierence between the pressures in the cylinder's upper and lower chambers. In all the three graphs, the enabling signal sent by the patient in order to initiate the step is represented by a dashed line. For the knee joint, a slight unforeseen ¯exion occurs during the stance phase, which in Fig. 9 corresponds roughly to the time periods 1±6.5 s and 9.5±15.5 s. This ¯exion takes place at the time the right limb is ¯exed, and is due to the type of pneumatic circuit used. It was not eliminated because it is in any case close to the physiological behaviour (Fig. 1).
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Fig. 9. Left knee angle.
In the graph shown in Fig. 10, which refers to the hip joint, the positive angles denote ¯exion while the negative angles denote extension. It is clear from the graph that hip and knee ¯exion are simultaneous; this shows that the operations preceding the swing phase are well coordinated, and minimizes energy expenditure. The torque curve in Fig. 11 has a number of signi®cant characteristics. First, it should be emphasized that this torque is equal to the drive torque acting on the joint only when ¯exion is greater than zero. During full extension, on the other hand, the actuator piston is at the end of stroke against the end plate, and part of the thrust force is unloaded against this constraint. The torque assumed in this situation, approximately ÿ60 Nm at 2 s, is maintained above the physiological value in order to prevent system collapse and provide a sucient margin of safety. Obviously, this torque can be varied as a function of supply pressure and in relation to the size of the pneumatic actuator used. When the right limb is ¯exed at 2.5 s, there is a reduction in absolute torque
Fig. 10. Left hip.
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corresponding to a slight ¯exion in Fig. 9, even though none of the left actuator valves are activated. The torque value is in any case sucient to guarantee stability, rising to the maximum value at approximately 6.5 s. Following the ¯exion command (at approximately 6.5 s), the absolute torque value drops, maintaining a modest positive value during the ¯exion movement. Similarly, it maintains a small negative value starting from the extension command at approximately 8 s, rising to the absolute maximum once full extension is reached. Using a force above the physiological value is generally dangerous, since in a paralyzed subject with no antagonist muscle force it could cause limb hyperextension with consequent ligament injury. In the case in question, this possibility does not in fact exist, as the cylinder is at the end of stroke at the time of maximum negative torque. A mechanical stop is thus provided which prevents 08 from being exceeded in extension. Subsequent tests indicated that even a small variation in the restrictors` ¯ow coecient resulted in dierent ¯ex/extension speeds, and hence in qualitatively dierent gaits. It was also observed that the absolute value of the extension moment drops when supply pressure is reduced, while the changes in the ¯exion torque are negligible. This is positive, as it is thus possible to come closer to the maxima and minima of the corresponding physiological values, and in particular to reduce the extension moment without causing appreciable changes in the ¯exion moment, which, as in the previous tests, remains within normal gait limits. 8.3. Tests with a paraplegic subject In experimental tests performed with a paraplegic subject having a T3 lesion level, the solution without hip actuation was found to be too dicult because of the increased weight of the knee actuators. By contrast, good results were achieved with the ``hybrid'' solution, where part of the energy is provided by the
Fig. 11. Torque trends on the knee joint.
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user and part by the pneumatic actuators. This solution calls for good coordination between the user and the phases of the gait cycle, which can be developed through training. An image sequence ®lmed during a test is shown in Fig. 12, while curves for the recorded magnitudes are given in the following ®gures. Tests were carried out with crutches, a walker and the parallel bars to maintain stability. As the parallel bars provide a greater sense of security, they were used in the majority of tests. It will be noted that the images in Fig. 12 coincide with the stick diagrams in Fig. 2 to a considerable extent. The cycle starts with left leg swing at approximately t = 45 s and ends at about t = 52 s. In particular, Fig. 12a (t = 45 s) shows the pre-swing of the left leg: this is still a double support phase, but body weight is almost supported by the stance leg, i.e., the right one. In Fig. 12b (t = 46 s), the swing leg passes the stance foot and the knee is still ¯exed, while the left hip angle is about zero. At t = 47 s (Fig. 12c), the left knee is completely extended, but the swing motion has not yet been accomplished, in fact the left hip angle is 108. The maximum stride length, corresponding to maximum hip ¯exion, is achieved at t = 48 s when the left foot touches the ground. Preparation for the swing phase of the right limb is depicted in Fig. 12e: the left arm is shifted ahead so that the body's center of gravity is also moved forward. This movement also causes the left foot tip to lower and the right heel to lift. Photographs of Fig. 12f±h, corresponding to t = 50, 51 and 52 s, respectively, are related to the swing phase of the right leg. The comparison between this experimental test and the one with the healthy subject shows a good level of correspondence between the dierent ¯exion angle curves. The slight dierences can be attributed to a number of factors, of which the most important is the fact that the orthosis was worn by a dierent user. For instance, the knee stance phase ¯exion (Fig. 13) which occurred in the earlier tests has now disappeared, and as a consequence the knee joint torque curve also diers (Fig. 14). This is due to the fact that the weight of the paraplegic patient is considerably lower than that of healthy user, and in this test the weight unloaded through the arms is probably higher. The dierence between left and right hip rotation range (Fig. 15), and hence stride length, is due to the user, whose left latissimus dorsi muscle injuries prevent him from exercising the same trunk extension and thus from making symmetrical strides. The maximum hip ¯exion and extension angles occur only when knee ¯exoextension is ®nished and the user extends the trunk and loads the rear cable. The values of the torque supplied by the pneumatic actuators are about 50% of the physiological value, as depicted in Fig. 16. In the test described here, activation of the hip and knee actuators is simultaneous. Dierent sequences of the valve command signals were used, but this pattern permitted a good gait and was also considered by the wearer to be the best.
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Fig. 12. Sequential images of a gait cycle.
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Fig. 13. Knee rotation angles.
Fig. 14. Hip rotation angles.
Fig. 15. Right knee torque.
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Fig. 16. Right hip torque.
9. Conclusions The orthosis presented herein is intended to assist locomotion in paraplegic subjects, including those with high-level lesions. In addition to being suitable for rehabilitation, it can be used on a daily basis in structured environments such as the home or workplace. The device is ideal for the rehabilitation stage, given the structure's modularity and the extremely ¯exible means used to regulate gait characteristics. Future work will address improvements in wearability and the gait cycle. As regards wearability, we intend to reduce the structure's weight and bulk, re-assess actuator placement and, if necessary, investigate other types of actuator. Changes in the gait cycle will focus on reducing physiological energy expenditure and will include the introduction of an ankle joint or a device capable of lifting the user during the one-legged stance phase so that the swinging limb will not strike the ground. Finally, a gait cycle which resembles the normal cycle in all respects, at least in the sagittal plane, can be achieved by using proportional interfaces for force control and for coordinating movements and the torques exerted by the actuators. Acknowledgements The research was supported by the Istituto Superiore della SanitaÁ. The authors would like to thank Steeper Europe BVBA for providing the ARGO. References [1] Nene AV, Hermens HJ, Zilvold G. Paraplegic locomotion: a review. Spinal Cord 1996;34:507±24.
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[2] Jasper J, Peeraer L, Van Petegem W, Van der Perre G. The use of an advanced reciprocating gait orthosis by paraplegic individuals: a follow-up study. Spinal Cord 1997;35:289±585. [3] Sykes L, Edwards J, Powell ES, Ross ER. The reciprocating gait orthosis: long-term usage patterns. Arch Phys Med Rehabil 1995;76:779±83. [4] Ogilvie C, Bower P, Rowley DI. The of paraplegic physiological bene®ts rthotically aided walking. Paraplegia 1993;31:111±5. [5] Stallard J, Major RE, Patrick JH. A review of the fundamental design problems of providing ambulation for paraplegic patients. Paraplegia 1989;27:70±5. [6] Lotta S, Fiocchi A, Giovannini R, Silvestrin R, Tesio L, Raschi A, Macchia L, Chiapatti V, Zambelli M, Tosi C, Baratta S, Franceschini M. Restoration of gait with orthosis in thoracic paraplegia: a multicentric investigation. Paraplegia 1994;32:608±15. [7] Glaser RM, Sawka MN, Wilde SW, Woodrow BK, Suryprasad AG. Energy cost and cardiopulmonary responses for wheelchair locomotion and walking on the tile and on carpet. Paraplegia 1981;19:220±6. [8] Massucci M, Brunetti G, Betti L, Franceschini M. Walking with the advanced reciprocating gait orthosis (ARGO) in thoracic paraplegic patients: energy expenditure and cardiorespiratory performances. Spinal Cord 1998;36:223±7. [9] Bernardi M, Canale I, Castellano V, Di Filippo L, Felici F, Marchetti M. The eciency of walking paraplegic patients using a reciprocating gait orthosis. Paraplegia 1995;33:409±15. [10] Solomov M, Shoji H, D'Ambrosia R, Douglas R. Electromechanical walking system for paraplegics. In: Proceedings of the IEEE 7th Annual Conference of the Engineering in Medicine and Biology Society. 1995. p. 4±7. [11] Vukobratovic M, Hristic D, Stojiljkovic Z. Development of active anthropomorphic exoscheletons. Medical and Biological Engineering 1974;12:66±80. [12] Rabischong P. Robotics for the Handicapped. In: Proceedings IFAC Control Aspects of Prosthetics and Orthotics (Columbus, OH, USA): May 7±9. 1982. p. 163±7. [13] Rabischong E, Sgarbi F, Rabischong P, Detriche JM, Pinguet N, Riwan A. Control and command of a six degree of freedom active electrical orthosis for paraplegic patient. In: Proceedings IEEE Inter. Workshop on Intelligent Robots and Systems, IROS '90 (Tsuchiura, Japan), July 3±6. 1990. p. 987±91. [14] Miyamoto H, Israel I, Miyamoto H, Mori S, Sand A, Sakurai Y. Approach to a powered orthosis for paralysed lower limbs. In: Proceedings ICAR '85. 1985. p. 458±551. [15] Miyamoto H. Powered orthosis for lower limbs Ð its structure and control. J Robotics and Mechatronics 1993;5(5):443±7. [16] Saunders JB, Inman VT, Eberhart HD. The major determinant in normal and pathological gait. The J of Bone and Joint Surgery 1953;35-A(3):543±58. [17] Beckett R, Chang K. An evaluation of the kinematics of gait by minimum energy. J Biomechanics 1968;1:147±59. [18] Belforte G, Sorli M, Gastaldi L. Applicazioni di dispositivi pneumatici al comando di tutori attivi. In: Proceedings Trasmissioni di Potenza (Assago, Italy): June 10±11. 1997. p. 118±31. [19] Belforte G, Sorli M, Gastaldi L. Active orthosis for rehabilitation and passive exercise. In: Proceedings Fourth International Conference on Simulation in Biomedicine, BIOMED '97 (Acquasparta, Italy): June 9±11. 1997. p. 199±208. [20] Belforte G, Gastaldi L, Sorli M, Manuello Bertetto A, Mazzeo M. Active orthosis: experiments with a healthy subject. In: Proceedings 8th International Workshop on Robotics, RAAD '99 (Munchen, Germany): June 17±19. 1999. p. 317±22. [21] Vaughan CL, Davis BL, O'Connor JC. Gait Lab. Champain, IL: Human Kinetics Publishers. [22] Steeper Advanced Reciprocating Gait Orthosis. Fabrication and Fitting Manual. Steeper. [23] Ferraresi C, Franco W, Manuello Bertetto A. Straight ®bbers pneumatic muscle: an actuator with high traction force. In: Proceedings 6th Scandinavian International Conference on Fluid Power, SICFP '99. (Tampere, Finland): May 26±28. 1999. p. 787±98. [24 Bouteille D, Belforte G. Automazione ¯essibile elettropneumatica e pneumatica. Milano, Italy: Tecniche Nuove, 1987.
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Glossary Anterior: Toward the front of the body (i.e., the toes are anterior to the heel) Extension: straightening movement of a joint, whereby the two contiguous segments form a more or less obtuse angle (for the hip joint, for example, extension brings the thigh posterior to the trunk) Flexion: bending movement of a joint, whereby the two contiguous segments form a more or less acute angle (for the hip joint, for example, ¯exion brings the thigh anterior to the trunk) Dorsi¯exion: lifting the tip of the foot or hand Plantar ¯exion: lowering the tip of the foot Inferior: with the body erect, the position toward the feet Hyperextension: extension beyond the anatomical position Lateral: farther from the medial plane Medial: toward the midline of the body, or on the median plane Frontal plane: plane dividing body into an anterior portion and a posterior portion. Also called coronal plane Sagittal plane: plane dividing the body into a right and left half Transverse plane: plane dividing the body into a superior portion and an inferior portion Posterior: towards the back of the body (i.e., the heel is posterior to the toes) Rotation: movement whereby a bone or a series of bones rotates around an axis passing through a joint Superior: with the body erect, toward the head Stance phase: gait cycle phase when the foot is on the ground Swing phase: gait cycle phase when the foot is lifted o the ground