PULSE WAVE VELOCITY DETERMINATION

Abstract

A system for determining a pulse wave velocity comprises a body lumen insertable device with an actuator for providing a pulse in the vicinity of a vessel wall and a sensor for sensing arrival of the pulse. A pulse wave velocity is obtained from the time difference between actuation of the actuator and sensing of the pulse. This system is used to create an artificial pulse which is transmitted along a vessel by a known distance before detection by a sensor.

Claims

1. A system for determining a pulse wave velocity of a pulse traveling along a wall of a vessel or organ, or a characteristic of the wall of the vessel or organ based on the pulse wave velocity, the system comprising: a body-implantable device which comprises: a cuff for implanting around the vessel or organ; an actuator for providing, through actuation of the actuator, a pulse in the vicinity of the vessel or organ wall at a first location along the vessel or organ; and a sensor for sensing the pulse at a second location along the vessel or organ, the second location being different from the first location; and a controller for: controlling the actuator and analyzing the sensed pulse; and determining, from a time difference between providing the pulse and sensing of the pulse, the pulse wave velocity or the characteristic of the vessel or organ wall based on the pulse wave velocity.

2. (canceled)

3. (canceled)

4. The system of claim 1, wherein the actuator is for compressing the vessel or organ to induce a pressure wave.

5. The system of claim 1, wherein the actuator is located at a point location or around a circumference of the vessel or organ wall.

6. The system of claim 1, wherein: the actuator and sensor are on the same body-implantable device; or the body-implantable device comprises a first separately implantable element for the actuator and a second separately implantable element for the sensor.

7. The system of claim 1, further comprising a second sensor located with the actuator and a second actuator located with the sensor, thereby to enable a velocity measurement in opposite directions.

8. The system of claim 7, wherein the actuator and the second sensor are a first multifunctional device, and wherein the sensor and the second actuator are a second multifunctional device.

9. The system of claim 1, wherein the actuator is adapted to provide a pulse to the vessel or organ wall for travelling along the vessel or organ wall to the sensor.

10. The system of claim 1, wherein the actuator is adapted to provide a pulse to the vessel or organ wall for travelling through contents of the vessel or organ to the sensor.

11. The system of claim 1, wherein the controller is further adapted to provide an indication of vessel or organ wall stiffness.

12. The system of claim 1, wherein the controller is further adapted to provide an indication of blood flow velocity.

13. The system of claim 1, wherein the actuator comprises an electroactive polymer actuator.

14. The system of claim 1, wherein the sensor comprises a pressure sensor, motion sensor, or strain gauge.

15. (canceled)

16. (canceled)

17. A method for determining a pulse wave velocity of a pulse traveling through a wall of a vessel or organ, the method comprising: implanting the system of claim 1 into a body: controlling an actuator to provide the pulse in the vicinity of the wall at a first location along the vessel or organ through actuation of the actuator; controlling a sensor to sense the pulse at a second location along the wall different from the first location; and analyzing the pulse to determine the pulse wave velocity from a time difference between providing of the pulse and sensing of the pulse.

18. (canceled)

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0046] Examples of the invention will now be described in detail with reference to the accompanying drawings, in which:

[0047] FIG. 1 shows a first example of a system for determining a pulse wave velocity;

[0048] FIG. 2 shows blood flow and arbitrary pulse propagation may be superposed;

[0049] FIG. 3 shows a second example of a system for determining a pulse wave velocity;

[0050] FIG. 4 shows an example in which the actuator is part of a first body insertable device and the sensor is part of a second body insertable device;

[0051] FIG. 5 shows a modification in which there is an actuator and a sensor at each end of the implant; and

[0052] FIG. 6 shows two levels of vessel blockage controlled by the actuator.

DETAILED DESCRIPTION OF THE EMBODIMENTS

[0053] The invention provides a system for determining a pulse wave velocity, comprising an body insertable device with an actuator for providing a pulse in the vicinity of a vessel wall and a sensor for sensing arrival of the pulse. A pulse wave velocity is obtained from the time difference between actuation of the actuator and sensing of the pulse. This system is used to create an artificial pulse which is transmitted along a vessel by a known distance before detection by a sensor.

[0054] The system can be used for investigating a wall of a vessel or organ of a human or animal subject. The system can be used to probe the wall of any organ in a human or animal body. However of particular interest are vessels and their walls. Important vessels are those of the blood circulatory system such as an artery or aorta. Other vessels can be probed as well. These could include those of the lymphatic system, those of the airways such as bronchia or those of the digestive system, such as the esophagus or gastrointestinal tract.

[0055] The invention provides a system and method which makes use of the concepts of pulse wave velocity (PWV) and pulse transit time (PTT), which each relate to propagation of the pulse pressure wave along the arterial tree. It is known that PTT and PWV can be used as a predictor of blood pressure.

[0056] PTT is defined as the propagation time of a blood pulse wave to travel a certain distance in an arterial segment. In practice often the combination of electrocardiography (ECG) and photoplethysmography (PPG) is employed to measure pulse propagation time because of the convenient way to monitor the ECG R-wave and pulse arrival moment in a distal artery with PPG.

[0057] Thus a known distance can be chosen between actuator and sensor, i.e. between location of generation of the pulse and location of sensing of the pulse. The distance needed may depend on the pulse transmission properties of the tissue in the vicinity of the vessel. For example, assuming a sensor output sample rate of about 10 kHz (should be achievable with a fairly simple controller), which correlates to a sample time interval of 100 microseconds, the distance a pulse wave may travel within an aorta (vessel) is around 700 micrometers, as a pulse wave velocity in such vessel may be around 7 (plus or minus 2) meters/second. Further assuming a 1% or better accuracy of measuring this distance, the full distance between location of providing the pulse and sensing of the pulse needs to be 70 mm.

[0058] A longer distance and/or higher sample rate will increase the accuracy of measurement.

[0059] One can optimize e.g. for one or both of these. So with a sample rate of in the kHz ranges a distance range could be larger than 2 cm, larger than 4 cm or larger than 6 cm. A useful range can be between 5 cm and 10 cm, which may correlate with current stent sizes used in surgery. It will be clear that other ranges of distance can be chosen and also that other sampling rates may be chosen. More elaborate discussion of distances also in relation to signal analysis is provided herein below.

[0060] The distance can be measured form first location (where the actuator provides a pulse) to the second location (where the sensor senses the pulse).

[0061] The pulse wave velocity is related to the intrinsic elasticity of the arterial wall. PWV is increased in stiffer arteries and, when measured over the aorta, is an independent predictor of cardiovascular morbidity and mortality. Given the predictive power of PWV, identifying strategies that prevent or reduce stiffening may be important in prevention of cardiovascular events. One view is that aortic stiffness occurs as a result of atherosclerosis along the aorta. However, there is little or no association between PWV and classical risk factors for atherosclerosis, other than age and blood pressure.

[0062] Furthermore, PWV does not increase during early stages of atherosclerosis, as measured by intima-media thickness and non-calcified atheroma, but it does increase in the presence of aortic calcification that occurs within advanced atherosclerotic plaque. Age-related broadening of pulse pressure is the major cause of age-related increase in prevalence of hypertension and has been attributed to arterial stiffening.

[0063] Arterial stiffness describes the reduced capability of an artery to expand and contract in response to pressure changes. Parameters that describe vessel stiffness include compliance and distensibility. Compliance (C) is a measure of volume change (V) in response to a change in blood pressure (P):

[00001]$C=\frac{.Math.V}{.Math.P}$

[0064] In a stiff vessel the volume change, and therefore compliance, is reduced for any given pressure change. However, compliance also relates to initial arterial volume because a smaller volume reduces compliance for any given elasticity of the arterial wall. Distensibility (D) is compliance relative to initial volume:

[00002]$D=\frac{.Math.V}{.Math..Math.P.Math.V}$

[0065] Therefore, it relates more closely to wall stiffness. The consequence of reduced compliance/distensibility is increased propagation velocity of the pressure pulse along the arterial tree, i.e. the pulse wave velocity, which relates to arterial distensibility by the Bramwell and Hill equation (with as blood density):

[00003]$\mathrm{PWV}=\sqrt{\frac{1}{.Math.D}}=\sqrt{\frac{V}{p}.Math.\frac{.Math.P}{.Math.V}}$

[0066] However, most often PWV is calculated by measuring the time taken for a pressure pulse to travel between two set points. Typically carotid and femoral pulse waves are obtained with pressure sensitive transducers attached over the carotid and femoral arteries. The pulse transit time (PTT) is the time interval (t) between the onset of the carotid and femoral pulse wave upstroke (known as the foot-to-foot method). The pulse wave travel distance (d), between the carotid and femoral pulse wave recording point, is measured over the body surface with a tape measure.

[00004]$\mathrm{PWV}=\frac{d}{.Math.t}=\frac{d}{t_2-t_1}$

[0067] Commonly used points are the carotid and femoral artery because they are superficial and easy to access. It covers the region over the aorta that exhibits greatest age-related stiffening. However such an approach only can average the PWV over a big distance.

[0068] In principle the PWV/PTT values may be obtained from points which are closer together. However, a problem is the imprecise measurement of the distance between the two pulse wave detections, which is typically measured on the skin. Furthermore, in certain cases, for example if a patient has very irregular pulses and especially pulse wave shapes, the time interval (t) cannot be determined very well or is not continuous. That leads to an incorrect measurement of the PTT and consequently the PWV. Accordingly also other medical diagnostics based on determining PTT/PWV, such as e.g. arterial stiffness measurements or plaque detection, lack precision.

[0069] FIG. 1 shows a first example of a system for determining a pulse wave velocity. It comprises body insertable device 10 in this case in the form of an implantable device, which comprises an actuator 12 for providing a pulse 14 in the vicinity of a vessel wall 16 and a sensor 18 for sensing arrival of the pulse at a different location to the actuator. This may for example be at a location downstream of the actuator 12 having regard to the direction of blood flow, although the sensing may be upstream, or indeed pulse timing measurement may be performed in both opposing directions, as discussed below. A controller 20,22 is provided for controlling the actuator 12 and analyzing the sensor signal. The controller may include some parts which are implanted as part of the device 10, such as control circuitry 20, and some parts 22 which are external to the body insertable device, such as a processor for processing the signals. There is wireless communication 24 between the controller parts. The controller 22, 24 determines a pulse wave velocity from the time difference between actuation of the actuator and sensing of the pulse.

[0070] FIG. 1 shows a cuff-like body insertable device 10 which is positioned around a vessel, in particular an artery 26.

[0071] In the example of FIG. 1, it is assumed that the pulse travels through the blood between the actuator and sensor. There are other examples where the pulse travels more predominantly through the vessel wall, described further below. However, in all cases, the invention provides a measurement of the PWV.

[0072] The cuff is soft, with a rigidity much lower than the arterial stiffness, and it is wound around the artery. The orientation of the cuff may be dependent on the blood flow direction so that it has to be implanted with the sensor downstream of the actuator with respect to the blood flow direction. Other examples may be insensitive to implantation orientation as described further below.

[0073] The actuator is for example an electroactive polymer actuator. The sensor may also be implemented as an electroactive polymer pressure or deformation sensor. The actuator and the sensor are clamped either around the whole circumference of the vessel or at one side or edge point. The actuator and sensor may include rigid parts to prevent lift off from the artery when activated. This could be a mass on to which the actuator or sensor is attached, a rigid narrow band around the artery to which the actuator or sensor is attached, or the actuator or sensor themselves may surround the artery.

[0074] By positioning the whole cuff at a suitable place in the body, e.g. close to a bone, such that the clamping mechanism is mechanically stabilized by the bone itself, lift off from the artery may also be prevented.

[0075] When the actuator is activated, there is a compression of the artery beneath the actuator, or if the body insertable device is applied within the artery, the actuator may knock against the inner wall of the vessel, or if directed to the blood the actuator may generate the pulse wave directly in the blood. This leads to the generation of an artificial pressure pulse wave in the medium (i.e. blood) in the artery and/or in the artery wall.

[0076] Transmission of the pressure pulse through the blood will be considered to be the dominant propagation path. After a short time, this pulse wave will be detected by the sensor, located at the opposite end of the cuff. By calculating the time difference between the excitation of the artificial pulse wave and the reception, the pulse transit time can be defined very precisely and the distance between actuator and sensor can be assumed to be known very precisely. Therefore an artificial PWV.sub.ap value can be calculated very reliably independent of the natural human heart beat (pulse wave).

[0077] For the sensor, a piezoelectric or a piezo-resistive pressure sensor may be used. However, other sensors which are sensitive to motion (e.g. accelerometer) or mechanical deflection (e.g. strain gauge) may be used as well. Other known sensors (e.g. MEMS sensor which change in capacitance or other impedance) may be used.

[0078] Since any damping or change in stiffness or in reflection of the arterial wall will have an impact on the pulse propagation, the device may be used to identify some critical arterial diseases like plaque growth, stenosis and aneurysm.

[0079] The control circuit 20 for example controls the actuator to deliver a short on/off pulse but in principle also any other pulse shape may be used to control the actuator. The overall controller 20 and 22 is also used to read out and process the sensor signal, to calculate the PTT.sub.ap and hence PWV.sub.ap values. These calculated values may then be transmitted or stored in a memory for later read out or else the raw data may be provided for external calculation.

[0080] Instead of a single pulse, also several pulses (or even a continuous signal with a certain pulse repetition frequency) may be applied. In such case additionally the phase shift between pulse generation and reception may be considered as a possible additional differentiator.

[0081] The connections between the electronics 20 and the sensor and actuator may be realized by electrically isolated flexible wires (or flex-foil) or there may be an electrically isolated part of the cuff itself functioning as the electrical connection lines. There may instead be a cover over the cuff functioning as a connection device to the electronics. The external reading of the data may be achieved by connecting a reading device to it or more conveniently by using state of the art wireless communication principles, which may be an active transceiver device or even more conveniently a passive (inductance coupling based) transceiver.

[0082] In addition to the determination of the PWV.sub.ap and PTT.sub.ap values, also information about the stiffness can be derived from the sensor signal. The stiffness of the artery not only has an influence on the PWV.sub.ap and PTT.sub.ap values but also on the signal amplitude and shape. For example, a soft tissue damps more of the incoming pulse wave than a stiff(er) tissue. Therefore in addition to the pure time-dependent behaviour, the amplitude and shape of the sensor signal may be taken into account to derive information about the stiffness (and accordingly plaque growth) or other mechanical parameter.

[0083] It may be desired to calibrate the body insertable device prior to the application. Therefore the pulse wave velocity of the artificially generated pulse may be determined by applying the body insertable device to an artificial tissue (e.g. an artificial artery) or even natural artery (human or animal) under conditions similar to the application later. For example, the artery may be filled with blood (or a similar liquid) without generating an active blood flow (to determine the steady state PWV.sub.ap/PTT.sub.ap) and/or at different blood flow velocities and/or different blood pressures. Furthermore, this calibration may include tests with different stiffness values or added material (representing artificial plaque) to the tissue (e.g. artificial artery) to know the change of PWV.sub.ap and PTT.sub.ap values prior to the application as precisely as possible.

[0084] This calibration also enables the determination of the (absolute) blood flow velocity. The time interval t.sub.1 is determined by the device. However, this time interval is based on the superposition of the pulse wave velocity of the artificial pulse and the blood flow velocity, which will be assumed to be in the same direction. Therefore it is necessary to separate the blood flow velocity from the artificial pulse wave velocity. Although it can be assumed that the artificially generated pulse wave propagates much faster than the velocity of the blood flow (and thus the influence of the blood flow may be negligible), it may be of additional benefit to determine the real blood flow velocity based on this calibration.

[0085] With reference to FIG. 2, the superposition of blood flow and arbitrary pulse propagation can be described as follows:

[00005]$v_{b.Math.f}+v_{a.Math.p}=\frac{.Math.d}{.Math.t_1}$

[0086] In the above, v.sub.bf is the velocity of the blood flow and v.sub.ap is the artificial pulse velocity, d being the distance between pulse generation and reception and t.sub.1 the measured time interval.

[0087] According to the calibration, the propagation velocity of the artificial pulse wave is known, the time interval t.sub.1 is measured and the distance d is known prior to the application. Accordingly the blood flow velocity can be calculated.

[0088] As mentioned above, the artificial pulse may travel also along the arterial wall. This propagation path may be considered for example when aneurysms need to be detected, since in such case the propagation path certainly will change.

[0089] FIG. 3 shows an arrangement for this purpose. The actuator 12 and sensor 18 are provided along a direct path along the vessel wall. The actuator knocks on the outer vessel wall such that a pulse wave preferably propagates on and through the vessel wall. In response to a change of the wall-stiffness (e.g. plaque) or the wall configuration (e.g. aneurism) the pulse propagation will change. This change can be detected by the sensor.

[0090] The examples above show the pulse wave generated in a spot/point-like approach, and the actuator knocks only on a small restricted region of the vessel. However in such a case there is a high risk of non-detectability of plaque, if the actuator is not well positioned. Therefore a longer actuator may be used which is totally wrapped around the vessel, such that when actuated, the whole artery is slightly compressed and accordingly the whole circumference of the artery will be approached and finally investigated.

[0091] The examples above show a single body insertable device.

[0092] FIG. 4 shows an example in which the actuator 12 is part of a first body insertable device 10a and the sensor 18 is part of a second body insertable device 10b. The actuator and the sensor may then be spatially separated and not part of the same body insertable device. Accordingly the impact of the body insertable device itself (i.e., cuff) on the variation in stiffness of the artery may be reduced but also larger distances may be covered. The actuator and sensor may be connected to the same electronics or they may have their own circuits 20a, 20b as shown.

[0093] FIG. 5 shows a modification in which there is an actuator and a sensor at each end of the body insertable device.

[0094] Thus, there is a second sensor 18 located with the actuator 12 and a second actuator 12 located with the sensor 18.

[0095] A first advantage of this configuration is the detection of the natural pulse wave (originated from the heart). This natural pulse wave will first initiate a sensor signal at sensor 18 followed by the generation of a signal at sensor 18, even without any artificial pulse generation. Accordingly the calculated time interval would then correspond to the PWV initiated by the heart. This may be determined in parallel to the arbitrarily generated pulse wave velocity, since the natural pulse wave and arbitrary one may differ in their form shape, by driving the actuator with a pre-defined driving form.

[0096] Another advantage of this configuration is to determine blood flow velocity. A first time interval t.sub.1 may be determined by actuating the actuator 12 and sensing with sensor 18. Then a second time interval t.sub.2 is determined by actuating the second actuator 12 and sensing with the second sensor 18.

[0097] The blood flow velocity may be determined because in one case the artificial pulse wave is detected while having the same direction as the blood flow (when actuator 12 is used) and in another case it is detected while having the opposite direction as the blood flow (when actuator 12 is used). Accordingly t.sub.1 will be different (shorter) to t.sub.2 and thus a new time interval t=t.sub.2t.sub.1 can be calculated, which relates to the blood flow velocity. Mathematically this can be described according to:

[00006]$v_{b.Math.f}-v_{a.Math.p}=\frac{.Math.d}{.Math.t_2}$

[0098] A new time interval, which is the difference between the two measured values can be defined:

t=t.sub.2t.sub.1

[0099] By combining these two equations, the blood flow velocity can be obtained:

[00007]$v_{b.Math.f}=+\sqrt{v_{\mathrm{ap}}.Math.\left(v_{\mathrm{ap}}+2.Math.\frac{.Math..Math.d}{.Math..Math.t}\right)}$

[0100] It is well known that an EAP device can be used as actuator and sensor, even in parallel. Superposed over the actuation signal, a high frequency oscillating signal may be provided and impedance measurement of the EAP can be determined. This impedance varies with applied mechanical load, or pressure or (change of) other environmental conditions.

[0101] Thus, one or more multipurpose EAP generators (for sensing and actuation) may be used, for example one at each end of the body insertable device, able to generate a pressure pulse and to measure a change in pressure.

[0102] The multipurpose EAP generator can also be used as a parallel actuator, as described above. Hence the actuator can be configured to partially close the vessel during blood flow or PWV measurements as shown in FIG. 6. FIG. 6 shows two levels of vessel blockage controlled by the actuator 12 and this may be achieved by applying a DC actuation voltage over which the measurement actuation (AC voltage) is superimposed. This has three advantages.

[0103] First, in-line calibration can be performed. By partially closing the vessel the flow characteristics can be determined as a function of the diameter. This can help in determining the extent of a stenosis later in the lifetime of the body insertable device.

[0104] Second, for PWV measurements of the blood vessel characteristics the vessel can be fully blocked by the EAP generator element to avoid interference of the heart beat with the measurement, because the (elastic) movement of the vessel walls due to the heart beat pulse is shut off for the time of the measurement.

[0105] Third, for blood flow measurements the vessel can be partially blocked by actuating the EAP generator. This enables a series of measurements at different actuation levels, simulating a (partial) blockage. A series of measurements under different known conditions can increase the accuracy of the blood flow measurement and also aid in stenosis detection.

[0106] The measured blood flow will only significantly change if the partial closure of the vessel by the actuator is more than a stenosis (in the vicinity of the actuator). By sweeping over the full DC actuation range and measuring the effect of the (partial) vessel closure on blood flow, a measure for the extent of the stenosis can be determined, and without having to perform a calibration.

[0107] The effect is also valid for pulse flow detection in the opposite direction (as described above).

[0108] The invention may be applied to many applications. Of main interest is personal health devices. Although the application described above is focussing on the detection of natural plaque (growth), other applications may be possible as well. Additional applications are the possible detection of plaque growth initiated by implanted other devices such as e.g. a by-pass. Often these types of artificial devices can induce plaque followed by clogging. Yet another possible application would be to detect the development and growth of aneurisms.

[0109] As discussed above, a controller performs the data processing. The controller can be implemented in numerous ways, with software and/or hardware, to perform the various functions required. A processor is one example of a controller which employs one or more microprocessors that may be programmed using software (e.g., microcode) to perform the required functions. A controller may however be implemented with or without employing a processor, and also may be implemented as a combination of dedicated hardware to perform some functions and a processor (e.g., one or more programmed microprocessors and associated circuitry) to perform other functions.

[0110] Examples of controller that may be employed in various embodiments of the present disclosure include, but are not limited to, conventional microprocessors, application specific integrated circuits (ASICs), and field-programmable gate arrays (FPGAs).

[0111] In various implementations, a processor or controller may be associated with one or more storage media such as volatile and non-volatile computer memory such as RAM, PROM, EPROM, and EEPROM. The storage media may be encoded with one or more programs that, when executed on one or more processors and/or controllers, perform the required functions. Various storage media may be fixed within a processor or controller or may be transportable, such that the one or more programs stored thereon can be loaded into a processor or controller.

[0112] As mentioned above, the actuator and optionally also the sensor may be implemented using an electroactive polymer (EAP) device. EAPs are an emerging class of materials within the field of electrically responsive materials. EAPs can work as sensors or actuators and can easily be manufactured into various shapes allowing easy integration into a large variety of systems.

[0113] Materials have been developed with characteristics such as actuation stress and strain which have improved significantly over the last ten years. Technology risks have been reduced to acceptable levels for product development so that EAPs are commercially and technically becoming of increasing interest. Advantages of EAPs include low power, small form factor, flexibility, noiseless operation, accuracy, the possibility of high resolution, fast response times, and cyclic actuation.

[0114] The improved performance and particular advantages of EAP material give rise to applicability to new applications. An EAP device can be used in any application in which a small amount of movement of a component or feature is desired, based on electric actuation. Similarly, the technology can be used for sensing small movements.

[0115] The use of EAPs enables functions which were not possible before, or offers a big advantage over common sensor/actuator solutions, due to the combination of a relatively large deformation and force in a small volume or thin form factor, compared to common actuators. EAPs also give noiseless operation, accurate electronic control, fast response, and a large range of possible actuation frequencies, such as 0-1 MHz, most typically below 20 kHz.

[0116] Devices using electroactive polymers can be subdivided into field-driven and ionic-driven materials.

[0117] Examples of field-driven EAPs include Piezoelectric polymers, Electrostrictive polymers (such as PVDF based relaxor polymers) and Dielectric Elastomers. Other examples include Electrostrictive Graft polymers, Electrostrictive paper, Electrets, Electroviscoelastic Elastomers and Liquid Crystal Elastomers.

[0118] Examples of ionic-driven EAPs are conjugated/conducting polymers, Ionic Polymer Metal Composites (IPMC) and carbon nanotubes (CNTs). Other examples include ionic polymer gels.

[0119] Field-driven EAPs are actuated by an electric field through direct electromechanical coupling. They usually require high fields (tens of megavolts per meter) but low currents. Polymer layers are usually thin to keep the driving voltage as low as possible.

[0120] Ionic EAPs are activated by an electrically induced transport of ions and/or solvent. They usually require low voltages but high currents. They require a liquid/gel electrolyte medium (although some material systems can also operate using solid electrolytes).

[0121] Both classes of EAP have multiple family members, each having their own advantages and disadvantages.

[0122] A first notable subclass of field driven EAPs are Piezoelectric and Electrostrictive polymers. While the electromechanical performance of traditional piezoelectric polymers is limited, a breakthrough in improving this performance has led to PVDF relaxor polymers, which show spontaneous electric polarization (field driven alignment). These materials can be pre-strained for improved performance in the strained direction (pre-strain leads to better molecular alignment). Normally, metal electrodes are used since strains usually are in the moderate regime (1-5%). Other types of electrodes (such as conducting polymers, carbon black based oils, gels or elastomers, etc.) can also be used. The electrodes can be continuous, or segmented.

[0123] Another subclass of interest of field driven EAPs is that of Dielectric Elastomers. A thin film of this material may be sandwiched between compliant electrodes, forming a parallel plate capacitor. In the case of dielectric elastomers, the Maxwell stress induced by the applied electric field results in a stress on the film, causing it to contract in thickness and expand in area. Strain performance is typically enlarged by pre-straining the elastomer (requiring a frame to hold the pre-strain). Strains can be considerable (10-300%). This also constrains the type of electrodes that can be used: for low and moderate strains, metal electrodes and conducting polymer electrodes can be considered, for the high-strain regime, carbon black based oils, gels or elastomers are typically used. The electrodes can be continuous, or segmented.

[0124] A first notable subclass of ionic EAPs is Ionic Polymer Metal Composites (IPMCs). IPMCs consist of a solvent swollen ion-exchange polymer membrane laminated between two thin metal or carbon based electrodes and requires the use of an electrolyte. Typical electrode materials are Pt, Gd, CNTs, CPs, Pd. Typical electrolytes are Li+ and Na+ water-based solutions. When a field is applied, cations typically travel to the cathode side together with water. This leads to reorganization of hydrophilic clusters and to polymer expansion. Strain in the cathode area leads to stress in rest of the polymer matrix resulting in bending towards the anode. Reversing the applied voltage inverts bending. Well known polymer membranes are Nafion and Flemion.

[0125] Another notable subclass of Ionic polymers is conjugated/conducting polymers. A conjugated polymer actuator typically consists of an electrolyte sandwiched by two layers of the conjugated polymer. The electrolyte is used to change oxidation state. When a potential is applied to the polymer through the electrolyte, electrons are added to or removed from the polymer, driving oxidation and reduction. Reduction results in contraction, oxidation in expansion.

[0126] In some cases, thin film electrodes are added when the polymer itself lacks sufficient conductivity (dimension-wise). The electrolyte can be a liquid, a gel or a solid material (i.e. complex of high molecular weight polymers and metal salts). Most common conjugated polymers are polypyrrole (PPy), Polyaniline (PANi) and polythiophene (PTh).

[0127] An actuator may also be formed of carbon nanotubes (CNTs), suspended in an electrolyte. The electrolyte forms a double layer with the nanotubes, allowing injection of charges. This double-layer charge injection is considered as the primary mechanism in CNT actuators. The CNT acts as an electrode capacitor with charge injected into the CNT, which is then balanced by an electrical double-layer formed by movement of electrolytes to the CNT surface. Changing the charge on the carbon atoms results in changes of CC bond length. As a result, expansion and contraction of single CNT can be observed.

[0128] Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims. In the claims, the word comprising does not exclude other elements or steps, and the indefinite article a or an does not exclude a plurality. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. Any reference signs in the claims should not be construed as limiting the scope.