Patent Publication Number: US-2020281531-A1

Title: Electromedical apparatus, a method for manufacturing the electromedical apparatus and a method for deploying the electromedical apparatus

Description:
TECHNICAL FIELD 
     The present invention relates to an electromedical apparatus, a method for manufacturing the electromedical apparatus and a method for deploying the electromedical apparatus, and particularly, although not exclusively, to an electromedical apparatus having a controllable mechanical structure. 
     BACKGROUND 
     Electromedical apparatus may be used to record electrical activity of a biological target. For example, electrocardiography may record the activity of the heart of a living species over a period of time, in which gel pads may be used to attach electrodes over the skin near near the position of the heart to detect the electrical changes on the skin that represent heartbeats. 
     Neural activities may also be recorded using electromedical apparatus, however the size of the electrodes of such apparatus designed for neural applications are much smaller when compared to electrocardiography, so as to precisely record activities of cells and tissues. These micro- or nano-scale electrodes may also be used in stimulating the cells in diagnostic tests or medical treatments. 
     SUMMARY OF THE INVENTION 
     In accordance with a first aspect of the present invention, there is provided an electromedical apparatus comprising: an electrical component arranged to facilitate an electrical signal transmission between an electrical device and a target; and a mechanical component physically connected to the electrical component; wherein the mechanical component includes a fluidic structure arranged to modify at least one physical property of the mechanical component, so as to facilitate a deployment of the electromedical apparatus on the target. 
     In an embodiment of the first aspect, the target is a biological target. 
     In an embodiment of the first aspect, the fluidic structure comprises at least one fluidic channel arranged to at least temporary accommodate a fluid therein. 
     In an embodiment of the first aspect, the mechanical component comprises a multi-layered structure including at least a first layer of material and a second layer of material. 
     In an embodiment of the first aspect, the at least one fluidic channel is defined by at least one trench formed on the second layer of material and encapsulated by the first layer of material. 
     In an embodiment of the first aspect, the first layer of material and the second layer of material are biocompatible, and are provided with desired mechanical and chemical properties, e.g. being mechanically flexible, electrically insulating, etc. 
     In an embodiment of the first aspect, the first layer of material includes polyimide. 
     In an embodiment of the first aspect, the second layer of material includes polydimethylsiloxane. 
     In an embodiment of the first aspect, the electrical component includes at least one metal electrode arranged on the first layer of material or the second layer of material. 
     In an embodiment of the first aspect, the at least one physical property includes a mechanical stiffness of the mechanical component, and wherein the mechanical stiffness is modified in response to a variation of an internal pressure of the fluidic channel. 
     In an embodiment of the first aspect, when the electrical apparatus operates in a first condition, the mechanical stiffness of the mechanical component is higher than that of the mechanical component when the electrical apparatus operates in a second condition. 
     In an embodiment of the first aspect, the mechanical component is arranged to prevent bending or buckling of the electromedical apparatus during the deployment of the electromedical apparatus on the target. 
     In an embodiment of the first aspect, the electrical component and the mechanical component combine to define an elongated body, and wherein the fluidic channel runs through at least a portion of the length of the elongated body. 
     In an embodiment of the first aspect, the fluidic structure further comprises an inlet and an outlet both in connection with the fluidic channel. 
     In an embodiment of the first aspect, the inlet and the outlet are defined at a first end of the elongated body. 
     In an embodiment of the first aspect, the electromedical apparatus further comprises a tapered tip defined at a second end of the elongated body. 
     In an embodiment of the first aspect, the fluidic channel comprises a U-shaped portion or a fluid returning portion of different shapes or designs at or proximate to the second end of the elongated body. 
     In an embodiment of the first aspect, the electrical component and the mechanical component combine to operate as a neural probe. 
     In an embodiment of the first aspect, the fluidic structure further comprises at least one fluidic cavity in fluid communication with a respective fluidic channel. 
     In an embodiment of the first aspect, the at least one physical property includes a physical dimension of each of the at least one fluidic cavity, and wherein the physical dimension is modified in response to a variation of an internal pressure in the fluidic channel and/or the fluidic cavity. 
     In an embodiment of the first aspect, in response to the variation of the internal pressure in the fluidic cavity, the at least one fluidic cavity is arranged to move a respective portion of the electrical component with respect to the target during the deployment of the electromedical apparatus on the target. 
     In an embodiment of the first aspect, the mechanical component is arranged to improve a physical contact between the electrical component and the target. 
     In an embodiment of the first aspect, the electrical component comprises an array of electrodes arranged to electrically contact with the target. 
     In an embodiment of the first aspect, the fluidic structure comprises a plurality of fluidic cavities, wherein the plurality of fluidic cavities defines an array of fluidic cavities. 
     In an embodiment of the first aspect, the size and dimension of the array of electrodes are different from those of the array of fluidic cavities. 
     In an embodiment of the first aspect, each of the electrodes in the electrical component aligns with a corresponding fluidic cavity in the mechanical component. 
     In an embodiment of the first aspect, the electrical component further comprises a counter electrode adjacent to each electrode in the array of electrodes. 
     In an embodiment of the first aspect, each electrode and counter electrode pair is arranged to estimate a position of the electrode with respect to the target so as to evaluate the physical contact between the electrical component and the target. 
     In an embodiment of the first aspect, the electrical component and the mechanical component combine to operate as a micro-electrode array apparatus. 
     In an embodiment of the first aspect, the electromedical apparatus further comprises a pump arranged to manipulate a flow of fluid through the fluidic channel under a predetermine pressure. 
     In accordance with a second aspect of the present invention, there is provided a method for deploying the electromedical apparatus in accordance with the first aspect on a target, comprising the steps of: increasing the internal pressure of the fluidic channel; implanting the neural probe in the target; and decreasing the internal pressure of the fluidic channel after a disposition of the electrical component in a first predetermined position. 
     In an embodiment of the second aspect, the method further comprises the step of controlling the internal pressure of the fluidic channel so as to move the electrical component from the first predetermined position to a second predetermined position. 
     In accordance with a third aspect of the present invention, there is provided a method for deploying the electromedical apparatus in accordance with the first aspect on a target, comprising the steps of: disposing the micro-electrode array apparatus on the target; an controlling the internal pressure of each of the plurality of fluidic cavities so as to improve the physical contact between the electrical component and the target at each respective portion of the electrical component. 
     In an embodiment of the third aspect, the method further comprises the step of estimating a position of each of the electrodes with respect to the target to evaluate the physical contact between each of the electrodes and the target. 
     In accordance with a fourth aspect of the present invention, there is provided a method for manufacturing an electromedical apparatus, comprising the steps of: defining a first layer of material with a fluidic structure; and combining the first layer of material with a second layer of material to form a multi-layer structure defining a mechanical component with the fluidic structure; and providing an electrical component to physically connect with the mechanical component, the electrical component is arranged to facilitate an electrical signal transmission between an electrical device and a target; wherein the fluidic structure is arranged to modify at least one physical property of the mechanical component, so as to facilitate a deployment of the electromedical apparatus on the target. 
     In an embodiment of the fourth aspect, the step of combining the first layer of material with a second layer of material comprises the step of stacking and binding the first layer of material onto a second layer of material. 
     In an embodiment of the fourth aspect, the first layer of material and the second layer of material are combined via chemical bondings. 
     In an embodiment of the fourth aspect, the first layer of material and the second layer of material are combined via a layer of adhesive material. 
     In an embodiment of the fourth aspect, the mechanical component is formed using a reversal imprint process. 
     In an embodiment of the fourth aspect, the step of defining a first layer of material with a fluidic structure comprises the step of applying the first layer of material on a stamp so as to construct the fluidic structure on the first layer of material. 
     In an embodiment of the fourth aspect, the step of defining a first layer of material with a fluidic structure comprises the step of coating a first anti-sticking layer on the stamp to modify a surface energy prior to applying the first layer of material on the stamp. 
     In an embodiment of the fourth aspect, the method further comprises the step of: transferring the first layer of material from the stamp to a handling substrate; 
     and releasing the first layer of material from the handling substrate after binding the first layer of material onto the second layer of material. 
     In an embodiment of the fourth aspect, the step of transferring the first layer of material from the stamp to a handling substrate comprises the steps of: coating a second anti-sticking layer on the handling substrate; stacking the first layer of material apply on the stamp to the handling substrate; and releasing the first layer of material from the stamp. 
     In an embodiment of the fourth aspect, the first or the second anti-sticking layer comprise FOTS and/or MOPTS, or other coating materials to provide the desired surface energy 
     In an embodiment of the fourth aspect, the step of providing the electrical component comprises the step of fabricating at least one electrode on the first and/or second layer of material. 
     In an embodiment of the fourth aspect, the step of fabricating the at least one electrode on the first and/or the second layer of material comprises the steps of: depositing a metal layer on the first layer of material and/or the second layer of material; and patterning the metal layer by photolighography and etching of the metal layer. 
     In an embodiment of the fourth aspect, the at least one electrode is fabricated prior to the step of combining the first layer of material with the second layer of material. 
     In an embodiment of the fourth aspect, the method further comprises the step of covering the at least one electrode with a passivation layer prior to combining the first layer of material with the second layer material. 
     In an embodiment of the fourth aspect, the method further comprises the step of exposing at least one electrical contact of the electrical component by etching. 
     In an embodiment of the fourth aspect, the at least one electrode comprises Au and/or Cr. 
     In an embodiment of the fourth aspect, the first layer of material and the second layer of material are biocompatible, and are provided with desired mechanical and chemical properties, e.g. being mechanically flexible, electrically insulating, etc. 
     In an embodiment of the fourth aspect, the first layer of material includes polyimide. 
     In an embodiment of the fourth aspect, the second layer of material includes polydimethylsiloxane. 
     In an embodiment of the fourth aspect, the first layer of material and the second layer of material are combine via a layer of binding material sandwiched between the first and the second layers of material. 
     In an embodiment of the fourth aspect, the binding material includes polymethyl methacrylate. 
     In an embodiment of the fourth aspect, the electromedical apparatus operates as a neural probe. 
     In an embodiment of the fourth aspect, the electromedical apparatus operates as a micro-electrode array apparatus. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee. 
       Embodiments of the present invention will now be described, by way of example, with reference to the accompanying drawings in which: 
         FIGS. 1A and 1B  are schematic diagraphs of an electromedical apparatus in accordance with embodiments of the present invention; 
         FIG. 2A  is a cross-sectional view of the electromedical apparatus of  FIG. 1A ; 
         FIG. 2B  is a cross-sectional view of a single electrode unit of the electromedical apparatus of  FIG. 1B ; 
         FIGS. 2C and 2D  are cross-sectional view of two embodiments of the electromedical apparatus of  FIG. 1B ; 
         FIG. 3  is an illustration showing a process for fabricating of an imprinting stamp for use in fabricating the electromedical apparatus in accordance with an embodiment of the present invention; 
         FIGS. 4A, 4B and 4C  are illustrations showing a process flow for fabricating the electromedical apparatus of  FIG. 1A ; 
         FIGS. 5A, 5B, and 5C  are illustrations showing a process flow for fabricating the electromedical apparatus of  FIG. 1B ; 
         FIG. 6  shows microscopic images showing the electrical component and the mechanical component of the electromedical apparatus of  FIG. 1A ; 
         FIG. 7  is an image showing a controller board in connection with the electromedical apparatus in accordance with embodiments of the present invention; 
         FIG. 8  is a microscopic image showing the mechanical component of the electromedical apparatus of  FIG. 1A ; 
         FIGS. 9A and 9B  are illustrations showing the parameters of the electromedical apparatus of  FIG. 1A  in an experiment for evaluating the buckling performance of the electromedical apparatus; 
         FIGS. 9C and 9D  are distribution graph and plot showing the results of the experiment according to  FIGS. 9A and 9B ; 
         FIGS. 10A, 10B, 10C, 10D, and 10E  are plots showing the relationships between critical buckling force vs internal fluidic pressure applied to the electromedical apparatus of  FIG. 1A  with different design parameters; 
         FIGS. 11A, 11B, and 11C  are plots showing measured and simulated critical buckling force at different fluidic pressure applied to the electromedical apparatus of  FIG. 1A  with different design parameters; 
         FIGS. 12A, 12B, and 12C  are plots showing measured critical buckling force at different fluidic pressure applied to the electromedical apparatus of  FIG. 1A  with different design parameters; 
         FIG. 13A  is an illustration showing a setup for gel movement test for the electromedical apparatus of  FIG. 1A ; 
         FIG. 13B  is a plot showing a relationship between the insertion force in the gel at different fluidic pressure applied to the electromedical apparatus of  FIG. 13A ; 
         FIGS. 13C and 13D  are micrographs and plots showing the displacement of the electromedical apparatus of  FIG. 13A  and comparison metal probes; 
         FIGS. 14A and 14B  are photographic images showing an insertion of the electromedical apparatus of  FIG. 1A  in a rat brain without and with an applied fluidic pressure in the fluidic channel; 
         FIG. 15A  is an illustration showing an adjustment of the position of the electrode in the electromedical apparatus of  FIG. 1B  towards the target of a retina surface with applied fluidic pressure; 
         FIGS. 15B, 15C and 15D  are a photographic image, a schematic diagram and a microscopic image showing the electromedical apparatus of  FIG. 1B ; 
         FIG. 16  is an image showing an implanted MEA in accordance with an embodiment of the present invention on a rat eye; and 
         FIGS. 17A, 17B and 17C  are plots showing the record signals in a stimulation test of the electromedical apparatus of  FIG. 1B  on a biological target with different stimulation current pulses. 
     
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
     The inventors have, through their own research, trials and experiments, devised that, electromedical apparatus, such as implantable neural probe(s) may be used to record and stimulate neural activities. For example, neural probes may be implanted into different organs or tissues (e.g. brain, eye, and ear). The neural probes include a biocompatibility for many neural probe applications, especially for those that require long term usage. 
     In some example embodiments, biocompatible polymer-based neural probes may be used for neural prosthetics. Preferably, different biocompatible polymers such as polydimethylsiloxane (PDMS), parylene-C, liquid crystal polymer (LCP), SU-8, and polyimide (PI) may be used to fabricate flexible neural probes to reduce human body responses after implantation. 
     However, these flexible polymer-based neural probes are relatively soft and therefore may cause difficulties in inserting into the targeted tissues. For example, due to the high flexibility of PDMS, these neural probes were difficult to be inserted into the brain or some other targeted tissues, and therefore may be more suitable to be used in surface neural prostheses applications. 
     In addition, the flexibility of the neural probes may also induce a certain degree of tissue damage during brain micro motions after the probes are implanted. Preferably, neural probes with higher flexibility may cause less tissue damage. However, a minimum probe stiffness may be needed to continue implant into the brain tissue. 
     In one example embodiment, flexible probe insertion may be accomplished using a probe stiffener, in which the flexible probe may be attached to a rigid silicon (Si) substrate with a biodegradable glue such as polyethylene glycol (PEG). Alternatively, biodissolvable materials such as silk polymer may be coated onto a surface of the neural probes to temporarily enhance neural probe stiffness during probe insertion. 
     In the abovementioned examples, the flexible probes are made more rigid during probe insertion, and then the probes are maintained to be highly flexible after the successful probe insertion. The process is irreversible. 
     In accordance with an embodiment of the present invention, the flexibility or the rigidity of the probes may be dynamically controlled. For example, the probes are provided with a fluidic channel in the probe body such that the flexibility of the neural probe may be dynamically controlled by manipulating the fluidic pressure in embedded channels. 
     Preferably, the neural probes may be controlled to be rigid during insertion, and then returned to inherent flexibility after the insertion. The neural probe flexibility may be adjusted by the using different probe materials, dimensions, and fluidic pressure in the embedded microchannels. 
     For example, neural probes may be fabricated in PDMS, PI, and/or PDMS/PI hybrid substrates. The polymer-based neural probes are biocompatible and their mechanical properties can be optimized. Preferably, PDMS and PI polymers have different Young&#39;s modulus, and therefore they can provide different degrees of stiffness. 
     Without wishing to be bound by theory, PDMS may be too soft for neural probe implantation into the brain tissue. PI is more rigid compared to PDMS. Accordingly, a combination of PDMS and PI layers in neural probes may be tuned to reach the appropriate flexibility. For neural probe implantation, the fluidic pressure in the embedded microchannels of PDMS/PI neural probes may be adjusted to provide a desired stiffness. The fluidic pressure in the microchannels could then be released after implantation so that the neural probes will be highly flexible for low tissue damage. 
     The inventors devised that neural probes in PDMS may be too flexible to be implanted into brain tissue, even with fluidic pressure in the embedded channels. On the other hand, the minimum critical buckling force for the PI probes was found to be 3.0 mN for 200 μm wide probes with 50 μm wide channels, which may be higher than the required force of 0.4-1.0 mN to penetrate into the brain tissue. 
     Without wishing to be bound by theory, hybrid PDMS/PI neural probes with the same dimensions and fluidic pressure may be more flexible with lower critical buckling force of 1.1 mN and rigid enough to penetrate the brain tissue. The flexibility of the PDMS/PI neural probes may be optimized at 0.25-1.25 mN depending on the fluidic pressure in the microchannel. 
     In one example embodiment, as fabricated neural probes were tested and inserted into a 0.6% agarose gel that mimicked the softness of the brain tissue. As the fluidic pressure was increased from 0 to 60 kPa, these probes were able to penetrate into the gel. After insertion, releasing the fluidic pressure allowed the probe to become more flexible to avoid tissue damage for chronic recording and stimulation. 
     With reference to  FIGS. 1A and 1B , there is shown an embodiment of an electromedical apparatus  100  comprising: an electrical component  102  arranged to facilitate an electrical signal transmission between an electrical device and a target  104 ; and a mechanical component  106  physically connected to the electrical component  102 ; wherein the mechanical component  106  includes a fluidic structure  108  arranged to modify at least one physical property of the mechanical component  106 , so as to facilitate a deployment of the electromedical apparatus  100  on the target  104 . 
     In this embodiment, the electromedical apparatus  100  comprises a plurality of electrodes or electrical contact pads which may transmit electrical signals between the target  104 , such as a neuron cell or a receptor in a biological target, and an external controller  110 . For example, control signal may be provided from the external controller  110  to stimulate the receptor cell in contact with corresponding one of the three contact pads of the electrodes  102 , or the controller  110  may pick up or detect a neural signal generated by a neuron cell in contact with the electrodes  102 . 
     Alternatively or additionally, other type of electrical or functional components such as sensors or electronic devices may be provided on the neural probe to provide functions other than signal transmissions. For example, temperature or pressure sensing electrodes may be provided to provide additional functions of the electromedical apparatus  100 . 
     Referring to  FIG. 1A , the electromedical apparatus  100  is a neural probe includes the electrodes  102  for transmitting electrical signals. The neural probe  100 A includes an elongated body  112 B (or a probe shank) and a probe tip  112 T at an end of the elongated body. The external devices such as a signal controller  110  may connect to the neural probe  110 A at the other end of the elongated body  112 B. 
     With reference also to  FIG. 2A , the neural probe  100 A comprises a mechanical component  106  which may be used for controlling a physical property, such as mechanical stiffness, of the neural probe  100 A. As discussed earlier, by increasing the mechanical stiffness of the neural probe  100 A, it may facilitate the deployment of the neural probe  100 A into a biological target  104  at a precise position and to prevent the probe to buckle during the implantation of the neural probe  100 A. 
     Preferably, the mechanical component  106  includes a fluidic structure  108 , such as a fluidic channel  108 A embedded in the probe body  112 B of the neural probe  100 A, for modifying the physical property of the probe. The position and dimension of the fluidic structure  108  are determined by the overall structure of the probe body  112 B and according to the application of the probe. For example, the fluidic channel  108 A may runs through a portion of the length of the elongated body  112 B which needs to have a dynamically controllable stiffness. 
     In this embodiment, the mechanical component  106  comprises a multi-layered structure including a first layer of material  202  and a second layer of material  204 , and the fluidic channel  108 A is formed by encapsulating the trenches formed on the second layer of material  204  with the first layer of material  202 . For example, a layer of polydimethylsiloxane (PDMS) may be patterned using an imprinting stamp to from the trenches and/or other three-dimensional structures, and then the PDMS layer  204  may be further bond to a layer of polyimide (PI)  202  using a PI/PDMS bonding method. Alternatively, the PDMS layer may be patterned using lithography. 
     As appreciated by a skilled person in the art, material other than PI or PDMS may be used to fabricate the electromedical apparatus  100 , e.g. polymers or other mechanically flexible material may be used to form the probe body  112 B. Preferably, the materials used are biocompatible for medical or biological applications. Besides, additional layers of material may be incorporated to form the multi-layer structure of the mechanical component  106  of the apparatus, which may further optimize the mechanical properties of the probe  100 A, or to provide additional structural/functional components on the probe  100 A based on different designs. 
     Referring back to  FIG. 1A , the fluidic structure  108  further comprises an inlet  108 X and an outlet  108 Y, provided at one end of the probe body  112 B, both in connection with the fluidic channel  108 A, such that fluid may be injected to the fluidic channel  108 A and is at least temporary held in the channel  108 A. In addition, the fluidic channel further comprises a U-shaped portion  108 U at or proximate to the second end, or adjacent to the probe tip  112 T, of the elongated body  112 B. In this configuration, injected fluids may run through the returning U-shaped path  108 U and are discharged at the outlet  108 Y which is provided at the same end of the inlet  108 X. Alternatively, the returning path may include other shape or design, for example it may be preferable in some embodiments that the electrical contacts/wires do not overlap with an underneath fluidic channel. 
     Preferably, a pump  114  may be used to manipulate a flow of fluid through the fluidic channel  108 A under a predetermine pressure. By continuously pumping or injecting a fluid into the fluidic channel  108 A and at the same time varying an internal pressure of the fluidic channel  108 A, the mechanical stiffness of the mechanical component  106  or the probe body  112 B may be dynamically controlled. Alternatively, the fluidic channel  108 A may be defined with a closed end in which fluids may be injected and extracted through the same inlet/outlet opening. 
     For example, the electromedical apparatus  100 , such as a neural probe  100 A referring to  FIGS. 1A and 2A , may operate in at least two conditions. In a first condition, the mechanical stiffness of the mechanical component  106  is higher than that of the mechanical component  106  when the neural probe  100 A operates in a second condition, in which the mechanical component  106  may prevent bending or buckling of the neural probe body  112 B during the deployment or implantation of the neural probe  100 A on the target  104 . 
     After disposing the neural probe  100 A in the desired position, the internal pressure in the channel  108 A may be decreased, such as by discharging the fluidic from the fluidic channel  108 A, thereby “returning” the probe  100 A to its flexible state. The internal pressure may be again increased or controlled by pumping in a fluid in the fluidic channel  108 A, to facilitate moving the implanted neural probe  100 A to a second position, without repeating the entire deployment process. 
     The soft neural probe body may reduce glial cell formation in the inserted region and prevent signal quality degrading. In some occasions, glial cells may form during stimulation or neural activity recording, the neural probe position may be adjusted by applying fluidic pressure inside the fluidic channel  108 A, thereby facilitating relocation of the probe  100 A to an area without damaged cells. 
     Referring to  FIGS. 1B, 2B and 2C , in an alternative embodiment, the electromedical apparatus  100  is a micro-electrode array (MEA) apparatus  100 B. For example, the MEA  100 B comprises an electrical component  102  having three electrodes arranged in a row, and a fluidic structure  108  comprises an array of fluidic cavities  108 B, with each of the fluidic cavities  108 B connected to a fluidic channel  108 A for transferring fluidic into and out of the fluidic cavities  108 B, e.g. with the help of an external pump  114  in communication with the fluidic channel  108 A. 
     The mechanical component  106  may be used improve a physical contact between the electrical component  102  and the target  104 , for example when the MEA  100 B is deployed to contact with a biological target  104  with non-planar surface such as on a retinal surface of an eye. The retinal surface may include a complex topography which is not perfectly fit with the curvature of a substrate of the MEA  108 B, which may results in gaps between the electrodes  102  and the retina cells, thereby causing bad contacts between the electrodes  102  and the retina at or around those regions with gaps between the two. 
     In this example, the physical property of the fluidic structure  108  includes a physical dimension, i.e. internal volume of each of the fluidic cavities  108 B. By injecting fluids into the internal cavities  108 B through the associate fluidic channels  108 A, the internal pressure and thus the internal volume of the cavities  108 B may be dynamically controlled. Preferably, the physical dimension is modified in response to a variation of an internal pressure in the fluidic channel  108 A and/or the fluidic cavity  108 B. 
     Referring to the cross sectional view of a portion of the MEA apparatus  100 B as shown  FIGS. 2B and 2C , the working principle of such control is similar to that of filling a balloon to increase the volume of the balloon, therefore eliminating or at least decreasing the size of the gap between the contact electrode  102  and the target  104 /retinal surface, and hence improving the physical contact with the target  104  by pushing the electrodes  102  towards the target  104 . 
     For example, after disposing the micro-electrode array apparatus  100 B on the target  104 , the operator may control the internal pressure of each of the plurality of fluidic cavities  108 B so as to improve the physical contact between the electrical component  102  and the target  104  at each respective portion of the electrical component  102 . 
     Preferably, each of the electrodes  102  in the electrical component  102  aligns with a corresponding fluidic cavity  108 B in the mechanical component  106 , so as to facilitate a pixel-to-pixel manipulation of the individual cavities. Alternatively, the two arrays may be offset or misaligned to each other, or the size/dimension of the two arrays may be different. 
     In response to the variation of the internal pressure in the fluidic cavity, the fluidic cavities  108 B may move a respective portion of the electrical component  102  with respect to the target during the deployment of the electromedical apparatus  100  on the target. Therefore, the two arrays in the electrical and the fluidic cavities  108 B are not necessarily the same. For example, referring to  FIG. 2D , a fluidic cavity  108 B with substantially bigger planar area may be responsible for improving a larger portion of the MEA  100 B covering multiple electrodes  102  may be defined, and therefore may require fewer individual fluidic channels  108 A in the apparatus. In addition, the shape of the fluidic cavities  108 B may not be the same and may include shapes other than round shape in different designs. 
     Optionally, the electrical component  102  further comprises a counter electrode  102 C adjacent to each electrode in the array of electrodes  102 . The counter electrode  102 C may be used for estimating a position of the electrode with respect to the target  104  so as to evaluate the physical contact between the electrical component  102  and the target  104 . For example, after disposing the micro-electrode array apparatus  100 B on the target  104 , by measuring an electrical resistance between the ring-shaped counter electrode  102 C and the round electrode  102 , a high resistance may represent a bad contact (due to a large gap) between the electrodes  102  to the target surface, then the corresponding fluidic cavities  108 B may be adjusted to push the electrodes  102  towards the target surface to improve the physical and hence the electrical contacts, and the quality of such contacts may be simultaneously monitored using the counter electrodes  102 C. 
     In addition, the ring-shaped counter electrodes  102 C may create an electrical isolating boundary during retinal stimulation to prevent cross talk of adjacent stimulating/main electrodes  102  due to electrical current leakage, for example by grounding all the counter electrodes  102 C during operation of the MEA  100 B. 
     Alternatively, the counter electrodes  102 C may be defined with different shape and dimension, provided that the counter electrode is disposed adjacent to the main electrode for signal transmission. 
     With reference to  FIG. 3 , there is shown an example embodiment of a fabrication of an imprinting stamp  300  for use in the subsequent fabrication of the fluidic structures  108  in the electromedical apparatus  100  such as the neural probes  100 A and the MEA apparatus  100 B as discussed above. 
     In this example, the imprinting stamp  300  may be used for patterning PDMS using a reversal imprinting method, in which PDMS material may be deposited on the stamp  300 , then the cured PDMS material may be detached and further transferred to another substrate or layer of material. 
     For example, silicon (Si) stamp  300  with 3D patterns of trenches or microchannels may be fabricated for reversal imprinting PDMS channel plate  204 , the PDMS layer may be stacked and chemically bonded on the PI cover plate  202  using 1 μm thick poly(methyl methacrylate) (PMMA) as an intermediate bonding layer. With such PMMA layer, chemical bondings form between he PI and PDMS layer and firmly combine the two layers of material. Alternatively, the first layer of material and the second layer of material are combined via a layer of adhesive material. 
     For the channel shape in Si stamp, 1 μm thick AZ nLOF 2070 (diluted) negative photoresist  302  may be spin-coated on Si at 3000 rpm for 60 s and soft baked at 130° C. for 60 s on a hot plate. The photoresist  302  may be patterned by UV exposure and post-baked at 130° C. for 60 s. After the negative photoresist is developed, a 5 μm thick AZ 6130 positive photoresist  304  may be spin-coated at 3000 rpm for 60 s and baked at 90° C. for 10 min. After patterning the positive photoresist  304 , the Si stamp  300  is ready to be dry etched. 
     A deep reactive ion etching (DRIE) (LPX ICP LE0729, SPTS) system may be used to etch the Si stamp in 2 steps. Initially, Si material of a thickness of 20 μm is etched to define the channel depth and the remaining negative photoresist  304  may be removed using O 2  plasma. The Si stamp may be further etched for an additional 40 μm to realize 60 μm total channel plate thickness. The remaining positive photoresist  302  may then be removed in acetone for 10 min. 
     With the Si stamp or mould  300 , the fabrication of the electromedical apparatus  100  may follow with the steps of defining a first layer of material  204  with a fluidic structure  108 ; combining the first layer of material  202  with a second layer of material  202  to form a multi-layer structure defining a mechanical component  106  with the fluidic structure  108 ; and providing an electrical component  102  to physically connect with the mechanical component  106 . 
     Using a reversal imprint process, the Si stamp  300  may be coated with trichloro(1H, 1H, 2H, 2H-perfluorooctyl) silane (FOTS) (Sigma-Aldrich, 97%), which acted as an anti-sticking layer  305 . A 60 μm thick 10:1 (weight ratio, w/w prepolymer to curing agent) PDMS  204  (Sylgard 184, Dow Corning) may be spin-coated at 1000 rpm for 60 s on the Si stamp  300  and reversely imprinted onto a handling substrate  306 , such as another Si substrate at 130° C. and pressure of 40 bar for 10 min. 
     With reference to  FIG. 4A , on this handling substrate  306 , the surface may be coated with 1:4 FOTS/methacryl oxypropylene trichlorosilane (MOPTS)  307  to provide higher surface energy than the FOTS coated Si stamp  300 , such that the PDMS layer coated on the stamp may be transferred to the handling substrate  306  when the stamp is released from the substrate after pressing the PDMS layer onto the substrate. 
     The ratio of FOTS and MOPTS may be tuned according to the desired surface energy on both the stamp and the handling substrate. Alternatively, other coating material may be used to modify the surface energy of the surfaces of the stamp or the substrates as desired. 
     Due to different surface energy between the Si stamp  300  and the Si handling substrate  306 , the PDMS-based channel plate  204  could be detached from the Si stamp  300 , and transferred to the Si handling substrate  306 , thereby providing a PDMS layer  204  with the fluidic structures  108  formed on it. 
     The fluidic channels  108 A are defined by encapsulating the trenches formed on the PDMS layer  204  and encapsulated with an additional layer of material, such as a layer of polyimide (PI)  202 . For example, referring to  FIG. 4B , the fabrication of a PI covering layer  202  may start with a clean glass or silicon substrate  308 . 1 μm PMMA  310  may be first spin-coated on the glass substrate at 3000 rpm and baked at 240° C. for 30 min to facilitate the final release of the neural probe  100 A from the glass substrate  308 . Then 15 μm PI  202  may be spin-coated at 1000 rpm for 60 s and baked at 180° C. for 5 h, followed by 240° C. for 4 h. 
     The electrical component  102  or the electrodes may be provided on the first or the second layer of material, based on different application of the electromedical apparatus  100 . For example, to fabricate a neural probe  100 A, the metal layer  312 , which may include multiple layers of Cr and Au, may be deposited on PI  202 , and the metal layer  312  may be patterned using lithography and etching method, and then the patterned metal layer  312  is subsequently covered by an additional layer of PI  202 . An additional 1 μm thick PMMA  310  may be spin-coated at 3000 rpm for 60 s as an intermediate bonding layer with the PDMS channel plate  204 . 
     Alternatively, other suitable metal, such as Pt may be used to define the electrodes or other electrical components of the electromedical apparatus  100 . 
     With reference to  FIG. 4C , PDMS channel plate  204  or the second layer of material may be treated by an O 2  plasma (Plasma-Term 790 Series) using 25 sccm O 2 , 65 W stage power, and 700 mTorr chamber pressure for 20 s. The PDMS channel plate  204  and the PI cover plate  202  may then be bonded together with low pressure for 2 min, followed by baking at 130° C. for 1 h. An inductively coupled plasma (ICP) system (Oxford Instruments, Plasmalab 80 Plus) may be used to dry etch 1/15 μm PMMA/PI at 10 sccm O 2 , 60 W stage power, 450 W Source power, and 10 mTorr chamber pressure for 15 min. PMMA and PI etch rate may be set at 1.0 and 0.8 μm/min, respectively, with the PDMS layer  204  as an etch mask. Finally, the entire PDMS/PI neural probe  100 A may be detached from the glass substrate  308 , for example by stripping away the PMMA layer  310  between the glass substrate  308  and the PI layer  202 , and lastly exposing the electrical contacts of the electrical component  102  by etching some PI covering the electrodes in the PI layer  202 . 
     With reference to  FIG. 5A  to  FIG. 5C , there is an alternative embodiment of the method for fabricating the electrical apparatus, such as an MEA apparatus  100 B, including the fluidic structure  108  and the metal electrodes  102 . A main different between the process for fabricating the MEA  100 B and the neural probe  100 A is that the electrodes  102  are provided on the PDMS layer  204  (second layer of material) in the MEA  100 B instead of on the PI layer  202  (first layer of material) in the neural probe  100 A. On the other hand, the reversal imprinting process, i.e. forming the fluidic structure  108  on the PDMS layer  204 , stacking and bonding of the two layers to complete the mechanical component  106  of the apparatus  100  are similar in two processes. 
     In this example, the process also starts with the preparation of the stamp  300 , preferable made of Si or SU-defining the structures of the microfluidic channels  108 A and the fluidic cavities  108 B. Firstly, referring to  FIG. 5A , silane  314  may be coated on the Si stamp  300 , and PDMS may be applied on the Si mould  300  to form the PDMS layer  204  with the fluidic cavities  108 B and the fluidic channels  108 A. 
     Secondly, the PDMS layer  204  on the Si stamp  300  (detached from a Si substrate with FOTS/MOPTS coated thereon) may be further deposited with a layer of metal  312  such as Cr, Au and/or Pt, followed by patterning the metal layer  312  by photolighography and etching of the metal layer  312  to define the metal electrodes  102 / 102 C on the PDMS surface. The patterned electrodes  102  may be further protected by a PDMS passivation layer followed by an electrode opening step using dry etching. Alternatively, the electrode opening step may be performed in the last step after combining the PDMS layer  204  with the PI layer  202  and releasing the combined layers from the glass handling substrates  308 . 
     Thirdly, referring to  FIG. 5C , the PDMS layer  204  with the 3D structure and the electrodes  102  may be transferred to a handling glass substrate  308  with PMMA coated thereon after detaching and releasing the PDMS layer  204  form the Si stamp  300 . Then the PDMS layer  204  on the handling glass substrate  308  may be flipped again on a PMMA/PI/PMMA/glass substrate followed by a PI/PDMS bonding process to stack and combine the PDMS 3D structure on the PI layer  202 . 
     Lastly, the MEA apparatus  100 B including the PDMS layer  204  and the PI layer  202  may be released from the two glass substrates  308  by stripping the PMMA layers coated on glass, and the PI layer  202  may be further etched using the PDMS layer  204  as the etching mask. 
     Apart from PDMS and PI, other materials may be used to fabricate the mechanical components in the electromedical apparatus. Preferably, the material is biocompatible, which may include most of the polymers. Alternatively, non-polymeric materials such as Si or quartz may be used if coated with biocompatible materials. Additionally or optionally, the first and second layers of material may include materials which may have other chemical or mechanical properties, such as being mechanically flexible and/or electrically insulating. 
     Preferably, one side (layer) of the probe is more flexible and the other side is more rigid. For the more flexible layer, for example, polyethene other than PDMS may be used. Polymers that contain long molecules that lie side by side can uncoil and slide past each other, making the material flexible. 
     For the more rigid layer of material, parylene and SU-8 other then PI could be used. Cross-linking is where the polymer chains are chemically joined together in places, by covalent bonds. The polymer molecules cannot slide over each other so easily. This makes materials tougher and less flexible, and they may not be easily stretched. 
     With reference to  FIG. 6 , there is shown a cross-sectional view of the neural probe  100 A and microscopic photographic images showing the mechanical and the electrical components  102  near the probe tip section. In this example, each of the electrical contact pads is round shape with a diameter of 30 μm, and the channel width of the fluidic channel  108 A is about 50 μm. 
     With reference to  FIG. 7 , there is shown a controller board  700  for use in connecting the electrodes  102  on the neural probe  100 A to an external controller  110 , a microfluidic inlet  702  is provided on the controller board for connecting the inlet  108 X and to the pump  114  for injecting the fluid into the fluidic channel  108 A for dynamically adjusting the internal pressure of the fluidic channel  108 A. 
     With reference to  FIG. 8 , there is shown a photographic image of an alternative example neural probe  100 A (without the electrodes) fabricated for further evaluation of the performance of the mechanical components  106  of the neural probe  100 A. 
     In the experiment, the embedded microchannels formed a ‘U’ shape in the neural probe near the probe tip, with single channel width of 12.5, 25, and 50 μm on each side of the U-shaped returning path, the total (in and out) channel width was 25, 50, and 100 μm, respectively. In addition, the width of the neural probes varied from 200 to 400 μm. It is devised that the channel width and fluidic pressure in the channels determined the flexibility of the neural probes, which could be optimized for implantation and reduced tissue damage. 
     The inventors has studied the mechanical performances of the fabricated neural probe, by measuring the flexibility of these neural probes using the buckling test that resembles the probe insertion into the targeted brain tissues. 
     In the experiment, neural probes with different materials and dimensions were simulated using the FEM-based COMSOL MULTIPHYSICS software. With reference to  FIGS. 9A and 9B , different neural probes were modeled in 3-dimensional Cartesian coordinate system with 2-dimensional views. The probe head was assumed to be fixed during the buckling. However, the probe tip was assumed to be sharp and could only move along the probe shank direction and could not slide along the surface. 
     With reference to  FIG. 9C , there is shown the the simulated von Mises stress distribution and the deflection of the probe along the probe shank. The critical buckling force was determined by measuring the force exerted by the neural probes on a microbalance (Sartorius BP 211D, Sartorius AG, Germany) with resolution of 0.01 mg. At the initial contact between the probe and the balance, the force value was set to zero. 
     With reference to  FIG. 9D , there is shown the changes of the axial force during the buckling test for a 60/15 μm thick PDMS/PI neural probe with probe width of 400 μm, channel width of 12.5 μm, and 120 kPa fluidic pressure in the microchannel. During the buckling test, after the probe tip contacted the microbalance surface, the probe shank gradually deformed as the probe was pushed down onto the microbalance by the induced force. At the threshold force of 1.9 mN, the probe buckled and the probe tip moved. All the critical buckling force was obtained based on the first buckling mode. 
     The following table shows a comparison of a neural probe in accordance with embodiments of the present invention with other example neural probes in controlling neural probes flexibility. It is shown that the neural probe of the present invention is only one that provided dynamic control of probe flexibility using adjustable fluidic pressure in the microchannels. In addition, the neural probes of the present invention may be designed to include different materials, such as PDMS and PI polymers, and with various probe dimensions. Advantageously, the neural probes may be optimized to the desirable low flexibility needed for implantation and high flexibility needed for in vivo operation with low tissue damage. 
     
       
         
           
               
               
               
               
               
               
               
             
               
                   
                   
               
               
                   
                   
                   
                 Channel 
                 Flexibility 
                 Buckling 
                 Insertion 
               
               
                   
                 Materials 
                 Probe Size 
                 Size 
                 Control 
                 Force 
                 Force 
               
               
                   
                   
               
             
            
               
                   
               
            
           
           
               
               
               
               
               
               
               
               
            
               
                 The 
                 PDMS/PI 
                 60 μm thick; 
                 12.5-50 
                 PDMS/PI; 
                 0.5-2.1 
                 mN 
                 0.8 mN 
               
               
                 present 
                   
                 200-400 μm 
                 μm wide; 
                 fluidic 
                   
                   
                 at 60 kPa 
               
               
                 invention 
                   
                 wide; 1 cm 
                 20 μm 
                 pressure in 
               
               
                   
                   
                 long 
                 deep 
                 channels 
               
               
                 Example 
                 Parylene-C 
                 24 μm thick; 
                 — 
                 Biodegradable 
                 2.6-300 
                 mN 
                 8-23 mN 
               
               
                 A 
                   
                 120 μm wide; 
                   
                 PEG and silk 
               
               
                   
                   
                 3 mm long 
               
               
                 Example 
                 Si/Parylene 
                 100 μm 
                 — 
                 Parylene tube 
                 540 
                 mN 
                 — 
               
               
                 B 
                 tube 
                 thick; 120 
                   
                 dimension and 
               
               
                   
                   
                 μm wide; 3 
                   
                 location 
               
               
                   
                   
                 mm long 
               
               
                 Example 
                 Si 
                 15 μm thick; 
                 — 
                 Breakable Si; 
                 1.7 
                 mN 
                 — 
               
               
                 C 
                   
                 100 μm wide; 
                   
                 biodegradable 
               
               
                   
                   
                 5.5 mm long 
                   
                 polymer 
               
               
                 Example 
                 Parylene 
                 20 μm thick; 
                 — 
                 Parylene 
                 0.5-1.3 
                 mN 
                 — 
               
               
                 D 
                   
                 200 μm wide; 
                   
                 stiffener 
               
               
                   
                   
                 2 mm long 
               
               
                 Example 
                 Parylene 
                 20 μm thick 
                 50-200 
                 Biodegradable 
                 12 
                 mN 
                 — 
               
               
                 E 
                   
                   
                 μm wide; 
                 PEG in 
               
               
                   
                   
                   
                 10 μm 
                 channel 
               
               
                   
                   
                   
                 deep 
               
               
                   
               
            
           
         
       
     
     With reference to  FIGS. 10A to 10C , there is shown the experimental results of critical buckling force for the PDMS-based neural probes with 200, 300, and 400 μm probe width, respectively, and  FIGS. 10D and 10E  illustrate the results for PI-based neural probes. In can be concluded that the critical buckling force for these PDMS-based probes were much smaller than the PI-based neural. 
     For 400 μm wide probes with 12.5 μm channel width, the critical buckling forces were found to be 0.06 and 7.30 mN for PDMS- and PI-based neural probes, respectively, with no fluidic pressure inside the channels. As channel width increased, the critical buckling force decreased for both types of probes. However, this effect was more significant for PI-based probes. The measurements also showed that the critical buckling force in the PDMS-based probes did not change significantly for probe widths varying from 200 to 400 μm due to PDMS&#39;s hyperelastic properties. Further increase of the fluidic pressure to 240 kPa inside the embedded channel would cause channel deformation and swelling since PDMS is relatively flexible when compared to PI. 
     However, the higher flexibility of the PDMS material that allows the deformation of the fluidic structure and hence the dimension of the body of the apparatus may be desired in the MEA apparatus. Details of the MEA apparatus will be discussed later in the disclosure. 
     Referring to  FIGS. 10D and 10E , the critical buckling forces for 60 μm thick PI-based neural probes with different probe width and channel width are shown in without and with 240 kPa fluidic pressure. The critical buckling force for these PI-based neural probes decreased with decreasing probe width or increasing channel width. Fluidic pressure inside the microchannel could increase the critical buckling force significantly. 
     For example, for 400 μm wide PI probe and 12.5 μm wide channel, the critical buckling force was 7.30 mN without fluidic pressure and it increased to 12.40 mN when 240 kPa fluidic pressure was applied. Similar increases were observed for 200 and 300 μm wide probes that the critical buckling force could be increased by applying fluidic pressure in the channels. 
     With reference to  FIGS. 11A to 11C , there is shown the experimental and simulation results of the critical buckling force for the 60/15 μm PDMS/PI neural probe with 12.5 μm channel width. For 200 μm wide neural probe with 12.5 μm wide channel, the critical buckling force was 0.29 mN with no fluidic pressure and it increased to 0.60 mN when the pressure was increased to 60 kPa. As the pressure was increased further to 120 kPa, the critical buckling force continued to increase to 1.17 mN. The measured critical buckling force for neural probes at different pressures closely matched with the simulated results. Similarly, referring to  FIGS. 11B and 11C , the measured and simulated results for 300 and 400 μm wide probes followed the same trends. 
     With reference to  FIGS. 12A and 12B , there is shown the effects of fluidic pressure on critical buckling force for neural probes with different channel widths and probe widths. The probes became stiffer with wider probes width, but the channel width had little effects on the critical buckling force. Channel width did not change the critical buckling force significantly because the probes consisted of 60 μm thick PDMS and only 10 or 15 μm thick PI layer with the channels. However, referring to  FIG. 12C , by changing the PI layer thickness from 10 to 15 μm the neural probes become stiffer. 
     With reference to  FIG. 13A , in another experiment, the fabricated probes were inserted into a 0.6% agarose gel that mimicked the brain tissue mechanical properties. With reference to  FIG. 13B , the plot illustrates measured insertion force as a function of fluidic pressure in the channels. The neural probes could be inserted into the gel when the applied force was large enough to shear through the agarose gel. 
     When no fluidic pressure was applied in the embedded microchannels, the probes were not able to penetrate into the gel. The insertion force was found to be 0.34, 0.51, and 0.63 mN for 200, 300, and 400 μm probe width, respectively, with no pressure inside the channels. The applied force deformed the surface of the gel without shearing through, therefore the probes were not able to penetrate into the gel. 
     Referring to  FIG. 13 , by increasing the fluidic pressure to 60 kPa, the probes may shear through the gel surface with insertion force of 0.45, 0.66, and 0.83 mN for 200, 300, and 400 μm probe width, respectively. Further increasing fluidic pressure to 120 kPa caused the probes to penetrate into the agarose gel. At higher pressure beyond 120 kPa, PDMS started to swell and became undesirable. 
     In order to characterize the flexibility of the neural probes, PI probes with width of 200, 300, and 400 μm were fabricated. These probes were inserted into the gel and the gel was moved laterally, while the probe displacements were tracked using a camera. Nickel/chromium (Ni/Cr) and copper (Cu) metal wire probes with a diameter of 70 and 150 μm, respectively, were used to compare the probes flexibility. Since a minimum force of 0.5 mN was needed to successfully implant the neural probes into the brain tissue. 
     As discussed earlier, the PDMS probes had critical buckling force below 0.5 mN and they were too soft to be inserted into the agarose gel. Therefore, the probe displacement measurements were performed in PI and metal wire probes. With reference to  FIG. 13C , the micrographs of the probe positions inside the gel at their starting and ending positions after the gel was moved laterally by 4 mm are shown, and the results are further illustrates in the plot as shown in  FIG. 13D , which shows the probe displacement, L, for different PI and metal wire probes. 
     Referring to  FIG. 13D , PI probes with width of 200, 300, and 400 μm had displacement that decreased with increasing probe width. In comparison, the displacement for the Ni/Cr and Cu wire probes were 3.21 and 0.35 mm, respectively. These results show that the polymer-based neural probes were more flexible than the metal wire probes, which have the advantages of inducing less tissue damage due to brain micro-motions. 
     With reference to  FIGS. 14A and 14B , there is shown another experimental result of deployment of neural probes in a rat brain without and with 120 kPa fluidic pressure respectively. It is obvious that the neural probe without a high fluidic pressure was not able to penetrate through the surface of the brain due to buckling. In contrast, the probe with 120 kPa fluidic pressure was successfully implanted. 
     Both the neural probe  100 A and MEA  100 B has embedded microfluidic channels  108 A where fluidic pressure can be applied. On the flexible neural probe, electrode  102  and connecting wires were patterned on the polyimide (PI)  202 , which is a non-stretchable material, only the overall probe rigidity is designed to be adjusted without deforming the metal layer. On the other hand, in MEA apparatus  100 B, metal layer  312  was patterned on PDMS  204  above the microcavity  108 B, referring to  FIG. 5C . When fluidic pressure is applied, due to the large membrane area to membrane thickness ratio in the cavity structure, the membrane can be dynamically inflated such that the position of electrodes  102  can be adjusted. 
     With reference to  FIG. 15A , there is shown an illustration of a cross-sectional view of the MEA apparatus  100 B similar to  FIG. 2B  with a single electrode unit  102  with a microfluidic cavity  108 B embedded underneath. There is a gap  120  between the electrode  102  and the retina surface  104 . By increasing the internal fluidic pressure in the cavity  108 B, the electrode  102  may be pushed towards the retina surface  104  to realize a physical contact between the electrode  102  and the retina cell  104 , thereby facilitating an electrical signal transmission between such electrode  102  and the contacted retinal cell(s)  104 . 
     By fine adjusting the positions of each of the electrode  102 , all the electrodes  102  can form a good physical contact with the target  104 , such as the retina cells, even though the substrate of MEA apparatus  100 B may not perfectly fit the topography on the retinal surface  104 . 
     Referring to  FIGS. 15B to 15D , there is shown two micrographs and a schematic drawing of a portion of the fabricated MEA apparatus  100 B and a single electrode unit  102  in the MEA apparatus  100 B in accordance with embodiments of the present invention. 
     With reference to  FIG. 16 , there is shown an MEA apparatus  100 B implanted in an eye  1600  of an adult male Sprague-Dawley rat, and signal transmission experiments were carried out. Stimulation of retinal cells was achieved by connecting the MEA on to a current source (200 μA, 600 μA, or 1000 μA) which generate current pulses of 0.2 ms duration and 30 s interval. The stimulating signal was delivered to the retina surface, while another neural probe or MEA implanted in the visual cortex in rat brain will pick up the transmitted neural signal. The results are illustrated in  FIGS. 17A to 17C . 
     These embodiments may be advantageous in that the electromedical apparatus such as neural probes and MEA apparatus may be dynamically manipulated with the utilization of embedded microchannels in the apparatus, so as to dynamically control their physical properties. 
     Advantageously, controlling the stiffness of neural probes may reduce motion induced tissue damage. By adjusting the polymer materials, neural probe dimensions, and fluidic pressure in the microchannels, probe stiffness could be dynamically controlled so that they could be inserted easily into the targeted tissue, but returned to be flexible after implantation by changing the fluidic pressure. 
     In addition, using the combination of PDMS and PI layers in the developed neural probes with embedded microchannels, the mechanical properties of the probes may be further optimized for deployment and becoming more flexible after penetration to avoid tissue damage. 
     Besides, the fluidic structure may be fabricated using different materials that allow an adjustment of dimension of it, such that the fluidic structure may be used in micromanipulation of the electrodes in the electromedical apparatus, such as individual electrodes in an MEA apparatus. This may further improve performance of retinal prosthesis or artificial vision apparatus. 
     It will be appreciated by persons skilled in the art that numerous variations and/or modifications may be made to the invention as shown in the specific embodiments without departing from the spirit or scope of the invention as broadly described. The present embodiments are, therefore, to be considered in all respects as illustrative and not restrictive. 
     Any reference to prior art contained herein is not to be taken as an admission that the information is common general knowledge, unless otherwise indicated.