Ординатура / Офтальмология / Английские материалы / Artificial Sight Basic Research, Biomedical Engineering, and Clinical Advances_Humayun, Weiland, Chader_2007
.pdf140 Liu et al.
achieve high power efficiency. To increase the Q of the coil, L and can be increased. Increasing the inductance L can be done by adding more turns to the coil. However, as the number of turns increases, the resistance increases as the equivalent wire length increases. Increasing the quality factor by increasing the frequency is limited by two factors. One is the self-resonating frequency of the coil. Another factor is that when the frequency is high, the eddy current in the wire induced by the magnetic field causes additional power dissipation and thus increases the equivalent resistance of the coil [21].
The design of a coil for power transmission involves the above factors in a non-trivial way. Given the physical size, the coil should be optimized through the right choice of diameter of wire, number of turns, number of strands, and operating frequency. Unfortunately, calculation of the inductance and resistance of a given coil involves magnetic field distribution and there are no simple analytical equations, which makes the optimization of the coils difficult [21]. A formal design procedure was developed in Ref. [12], which formulates a semianalytical relationship between the inductance and the resistance of a coil and finds an optimal coil structure within a given size limitation to achieve high power transfer efficiency.
Power Transmission and Recovery
The transmitting coil needs alternating current to generate an alternating magnetic field so that the required voltage can be induced across the receiving coil. This is done by a power amplifier that converts the DC energy from the battery to AC and drives the transmitting coil. Class-E power amplifier is a good choice due to its high power conversion efficiency (90–97%) and ability to generate high-amplitude output signals across the inductor while operating at a small supply voltage. The power coupled to the receiving coil is AC and needs to be converted to DC for the operation of implant electronics. As mentioned before, for biphasic stimulus of large number of electrodes, stimulation through two supply voltages is preferred. To convert the AC voltage into dual DC supply, two forms of rectification can be employed, with one favoring the power efficiency and the other favoring the size of the implant.
As shown in Figure 7.7a, the diode D1 and capacitor CS rectifies the AC voltage V2 to DC voltage VS. The series regulator provides the first stage voltage regulation followed by a shunt regulator. The shunt regulator provides the required supply voltages Vdd (positive) and Vss (negative) with a common ground (zero potential). However, for biphasic stimulation, the anodic stimulation does not occur at the same time as the cathodic stimulation. In this case, the anodic current ia flows through the electrode and the cathodic shunt regulator, as shown in Figure 7.7a. This results in additional power loss in the cathodic shunt regulator. The power dissipated on the cathodic shunt regulator Vss · ia is wasted, which is about half the total power delivered by the series regulator
Vdd + Vss · ia . In other words, if the required stimulation power is Pload, the series regulator needs to supply twice the power of that 2Pload .
7. Challenges in Realizing a Chronic High-Resolution Retinal Prosthesis |
141 |
Figure 7.7. Rectification Topologies (a) single diode, (b) dual diode.
The power dissipation in the cathodic shunt regulator can be avoided if a return path is provided for the anodic current, bypassing the cathodic regulator. The rectification topology shown in Figure 7.7b provides this return path. The two voltages are generated through a dual-diode rectification with the series regulators without any shunt regulators. This topology provides a low impedance return path for the stimulation current to the storage capacitors. Thus even when only anodic/cathodic stimulation is activated the current will not flow through the cathodic/anodic regulator as in the case of Figure 7.7a. This allows the regulated power to stimulate the tissue without any additional loss. However, this topology requires an additional diode and a capacitor compared to the topology in Figure 7.7a, which increases the size of the implant especially when off-chip components are used.
The choice between the two rectification topologies has to be made from the system design perspective. If the stimulation of different electrodes can be scheduled in a way such that at a given time, the combined anodic current of one set of stimulated electrodes is equal to the cathodic current of the rest of the stimulated electrodes, then the current passing through shunt regulators can be reduced. However, at the time of this writing, the effective stimulation pattern is not known. Thus the dual-diode topology is preferred to maintain the power efficiency.
142 Liu et al.
Choice of Storage Capacitor
The capacitor(s) CS shown in Figure 7.7 stores the charge that is to be delivered to the tissue through electrodes. It also acts as a filter to remove the ripple from the output of the diode. Due to the varying magnitude of the stimulus current, the capacitor needs to be large enough to store charge to avoid unacceptable voltage drop when charge is delivered to the electrodes. Figure 7.8 illustrates this problem for the dual-diode topology case. For the single diode topology, the analysis is similar but the requirement of the capacitance is even higher since the capacitor has to supply both anodic and cathodic currents. If N is number of electrodes stimulated, the total charge delivered to the tissue by the stimulator
Qstim = Istim · N · PWstim. The average current from source is Is_avg = Qstim/Tstim. The voltage drop Vcap during the stimulation can be calculated as Vcap =
1/CS · Istim · N − Is_avg PWstim. If N = 1000 Istim = 100 A PWstim = 1
millisecond, Tstim = 16 milliseconds and the maximum voltage drop the system can tolerate is 1 V, then the required capacitance is 93 75 F. This is far too large a capacitance to be implemented in an IC. Even an off-chip capacitor of this size will consume a large area for an implant. Fortunately, this is an extreme condition, when all the 1000 electrodes are stimulated at the same time, which is unlikely. In addition, if the stimulation of electrodes can be
Figure 7.8. Storage capacitor and related waveforms.
7. Challenges in Realizing a Chronic High-Resolution Retinal Prosthesis |
143 |
scheduled in a sequence such that the current consumption of the electronics does not vary much over time, the capacitance can be reduced. However, there is not enough knowledge at this point, about the stimulation pattern and the interactions between electrodes. While clinical experiments should focus on obtaining more insights on these parameters, the test device for these experiments should be designed with enough flexibility to facilitate the experiments. So, the capacitance should be chosen as large as possible which can be integrated within the available area.
Coil Coupling and Loading Variations
For the patient’s comfort, the external unit should have some freedom of movement instead of a rigid fixture, which may cause the change of distance and orientation angle between the transmitter and the receiver coils. As a result, the coupling between the two coils varies and the coupling coefficient is estimated to change from 0.08 to 0.24 for a distance variation from 15 to 5 mm. Furthermore, the power consumption in the implant unit may change significantly with stimulation data pattern through changes in stimulation currents. For example, during standby operation, less power is needed. Both the coupling and the loading variations may cause one of the following: (1) increased implant heating due to excess transmitted power, causing a detrimental effect on the tissue over long term; (2) decreased implant supply voltage/current due to insufficient transmitted power, causing improper device operation or shutdown. This poses challenges to electronic design to transmit the optimal power required by the implant.
Research efforts have been directed for compensating the loading/coupling variations. The work in Ref. [22] proposed to estimate the coupling coefficient by detecting the primary coil voltage and current. This method somewhat reduces the impact of coupling change, although it is only applicable to systems with constant loading. Furthermore, without knowing the exact information on received power at the secondary side, it cannot achieve optimal compensation. In Ref. [23], we implemented an adaptive wireless power telemetry system where the transmission power is controlled according to the needs of the implant. By detecting the power level inside the implant and wirelessly transmitting this information back to the external unit, a closed loop control was realized to achieve immunity to coupling and/or loading variations. This system requires reverse telemetry, communication from the implant to external unit, which will be discussed in the next section.
Wireless Data Communication
The data communication between the external unit and the retinal implant needs to be bi-directional. The one from external unit to implant is termed “forward data telemetry” and the one from implant to external unit is termed “reverse data telemetry.” Forward data telemetry transfers the parameters of the stimulation
144 Liu et al.
(converted from image information by signal processing unit) to the implant. Besides the image information, forward data telemetry also needs to relay configuration information to the implant. Reverse data telemetry transfers information to the external unit from the implant. The information includes implant status such as temperature and pH, received power level, impedance measurement data, etc.
Forward Data Telemetry
To achieve maximum flexibility of stimulation, forward data rate needs to be as high as possible. For high-density (1000 pixels) retinal implant, the estimated forward data rate is about 1 to 2 Mbps, as mentioned earlier. There are several ways to transmit information to the implant. The simplest method of data transmission to and from the implant is a percutaneous wire connecting the two units. The disadvantages of using a tethering wire as discussed in the power telemetry section apply here as well. Optical telemetry is another option for transmitting data, where a modulated beam of infrared light is used to encode information. This information would then be decoded at the receiving end using optoelectronic detectors and further processed using electronics. However, optical transmission efficiency decreases rapidly with skin thickness and it is more sensitive to relative dislocation between the transmission element and the receiving element compared to magnetic coupling with coils [24] as well as the ambient light sources. The challenges associated with inductive method of data telemetry will be discussed below.
To transmit data into implant, one way is to utilize existing power transmission link by modulating the data onto the power carrier. Another way is to use an RF link with frequency different from the power carrier. In the first case, the power carrier is modulated with the data, thus employing the power carrier as the data carrier. This type of telemetry system is termed ‘single band telemetry’ referring to the fact that the same carrier frequency is used for transmitting both power and data through the same physical means. The method has the obvious advantage of not needing a dedicated coil pair for data telemetry. However, it faces the challenge of high data rate required by high-density stimulation. Normally the power carrier frequency is less than 10 MHz to penetrate the human body without significant absorption and achieve high efficiency of power transmission. Achieving high data rates (1 to 2 Mbps) at a relatively low carrier frequency is a great engineering challenge. Another approach to high data rate and high efficiency power transmission, termed ‘dual-band telemetry’, has been proposed in Ref. [25]. Shown in Figure 7.9, the power and data are transmitted through separate pairs of coils using two different carrier frequencies, forming a hybrid dual-frequency link which allows optimization of efficiency power telemetry and data rate of the forward telemetry through allocation of different frequencies. The power carrier frequency fL is lower than the data carrier frequency fH.
7. Challenges in Realizing a Chronic High-Resolution Retinal Prosthesis |
145 |
Figure 7.9. Dual-band telemetry.
However, the dual-band approach faces the challenge of interference between the power and the data link. As shown in Figure 7.10, the four coils have six coupling coefficients, which make the design complicated. The power transmitter generates a strong magnetic field in order to transmit power to the implant. This magnetic field can generate a fairly high voltage on the data receiver coil, which may corrupt the data signal completely. One way to tackle this problem is to filter out the power interference before passing the data signal to the recovery circuits. This filter may consume additional power in the implant, which is not desirable. Another way is to minimize the coupling by relative positioning of the power and data coils. For example, the data coils L1_data and L2_data can be
Figure 7.10. Interference between power and data telemetry.
146 Liu et al.
placed orthogonal to the power coils L1_pwr and L2_pwr , which ideally reduce the coupling coefficients of K11 and K22 to zero. However, winding L2_data around
L2_pwr (to make the zero coupling of K22) creates a bump on coil L2_pwr, which is undesirable from the surgical perspective.
Compared to the single-band approach, the dual-band approach requires one more coil to be implanted, which increases the size of the implant. On the other hand, it can save power since it allows power link to be optimized without sacrificing for data transmission. Design of a dual-band telemetry system should try to minimize the size increase due to the extra coil.
Reverse Data Telemetry Design
Reverse telemetry may use another frequency as data carrier, similar to forward data telemetry. This approach is called ‘active reverse telemetry’. The data carrier may use the existing coil pair or another coil pair as physical link. In general, this approach requires an active driver to drive the coil/antenna. Another approach is called “passive reverse telemetry,” as it does not require a dedicated data driver. Figure 7.11 shows the principle of using passive telemetry to transmit data from the implant to the external unit. It should be noted that in a dual-band telemetry system, the reverse telemetry can use the power telemetry as its data carrier, since it does not require high data rate and the low frequency power carrier is a good candidate. As shown in Figure 7.11, the coil is not driven by any power amplifier. Instead, a change in the loading of the coil is introduced by turning on and off the switch S. This loading change at the secondary coil creates a detectable change in the coil current (hence called load shift keying [22]). A receiver in the external unit connected to the primary coil detects this change and recovers the data. The load shift keying shown in Figure 7.11 may generate a high voltage stress on the switch and the diode, when the switch is open, since interrupting current flow through an inductor produces a large voltage to counteract the change. This high voltage in turn makes it difficult to implement the switch in the IC, if it is above the operating voltage of the chosen process technology, and may require discrete components, which increases the size of the implant. Thus the reverse telemetry should be designed to maintain the advantage of low power dissipation while minimizing the discrete components and consuming as less space as possible.
Figure 7.11. Passive reverse telemetry.
7. Challenges in Realizing a Chronic High-Resolution Retinal Prosthesis |
147 |
Conclusions
We have examined several challenges in the electronics design toward the implementation of chronic high-resolution retinal prosthesis. The key challenges are to reduce power consumption, provide more flexibility, and reduce the size. As seen through the examples in the above sections, solving these challenges often involve direct tradeoffs among themselves. The stimulator circuit faces the challenge of providing maximum flexibility, which increases the power and area. The power telemetry faces the challenge of increase in area in the form of off-chip components like coil, diode and capacitor, while maintaining a high efficiency and long battery life. The data telemetry faces the challenge of supporting high data rate for maximum flexibility hence requiring an additional coil pair. A system design is highly critical in realizing a chronic high-resolution retinal prosthesis, during which the above tradeoffs should be considered to make the suitable choices. In addition to the design of the electronics, the choice of technology for implementing the circuits deserves careful attention. A small feature size technology with more metal layers is preferred to reduce the size of the IC. This in turn requires the stimulation voltages to be small enough to allow the use of such a process technology. This also poses challenge for circuit design to accommodate high voltages. Passive components such as the storage capacitor, coils, and active components such as the back telemetry switch are difficult to implement in an IC. Off-chip versions of these components, which can lend themselves to be integrated with the ICs through post-processing on the surface of the IC, are highly preferred over components that require to be connected using wires.
References
1.Margalit E and Sadda SR (2003) Retinal and optic nerve diseases. Artificial Organs 27:963–974
2.Humayun MS, Prince M, De Juan Jr E, Barron Y, Moskowitz M, Klock IB, and Milam AH (1999) Morphometric analysis of the extramacular retina from postmortem eyes with retinitis pigmentosa. Investigative Ophthalmology and Visual Sciences 40:143–148
3.Humayun MS, De Juan Jr E, Weiland JD, Dagnelie G, Katona S, Greenberg R, and Suzuki S (1999) Pattern electrical stimulation of the human retina. Vision Research 39:2569–2576
4.Humayun MS, Greenberg RJ, Mech BV, Yanai D, Mahadevappa M, van Boemel G, Fujii GY, Weiland JD, and De Juan Jr E (2003) Chronically implanted intraocular retinal prosthesis in two blind subjects. In: Proceedings of ARVO Annual Meeting, April 2003.
5.Weiland JD, Yanai D, Mahadevappa M, Williamson R, Mech BV, Fujii GY, Little J, Greenberg RJ, De Juan Jr E, and Humayun MS (2004) Visual task performance in blind humans with retinal prosthetic implants. In: Proceedings of the 26th Annual International Conference of the IEEE EMBS
148 Liu et al.
6.Weiland JD, Liu W, and Humayun MS (2005) Retinal prosthesis. Annual Review of Biomedical Engineering 7:361–401
7.Cha K, Horch K, and Normann RA (1992) Simulation of a phosphene-based visual field: visual acuity in a pixelized vision system. Annals of Biomedical Engineering 20:439–449
8.Cha K, Horch KW, Normann RA, and Boman DK (1992) Reading speed with a pixelized vision system. Journal of Optical Society of America A: Optics, Image Science and Vision 9:673–677
9.Cha K, Horch KW, and Normann RA (1992) Mobility performance with a pixelized vision system. Vision Research 32:1367–1372
10.Hayes JS, Yin JT, Piyathaisere DV, Weiland JD, Humayun MS, and Dagnelie G (2003) Visually guided performance of simple tasks using simulated prosthetic vision. Artificial Organs 27:1016–1028
11.Liu W and Humayun MS (2004) Retinal prosthesis. In: IEEE International SolidState Circuits Conference Digest of Technical Papers, pp 218–219
12. Kendir GA, |
Liu W, Wang G, Sivaprakasam M, Bashirullah |
R, |
Humayun |
MS, and Weiland JD (2005) An optimal design methodology |
for |
inductive |
|
power link |
with class-E amplifier. IEEE Transactions on |
Circuits and |
|
Systems – I 52:857–866 |
|
|
|
13.Wang G, Liu W, Bashirullah R, Sivaprakasam M, Kendir GA, Ji Y, Humayun MS, and Weiland JD (2004) A closed loop transcutaneous power transfer system for implantable devices with enhanced stability. In: Proceedings of the IEEE International Symposium on Circuits and Systems. pp 17–20
14.Sivaprakasam M, Liu W, Humayun MS, and Weiland JD (2005) A variable range biphasic current stimulus driver circuitry for an implantable retinal prosthetic device. IEEE Journal of Solid-State Circuits 41:763–771
15.Fink W, Tarbell M, Weiland JD, and Humayun MS (2004) DORA: Digital Object Recognition Audio-Assistant for the visually impaired. Investigative Ophthalmology and Visual Science 45:4201
16.http://www.mosis.org/products/fab/vendors/
17.Mahadevappa M, Weiland JD, Yanai D, Fine I, Greenberg RJ, and Humayun MS (2005) Perceptual thresholds and electrode impedance in 3 retinal prosthesis subjects. IEEE Transactions on Neural Systems and Rehabilitation Engineering 13:201–206
18.Soykan O (2002) Power sources for implantable medical devices. Medical Device Manufacturing and Technology
19.Grover FW (1973) Inductance Calculation Working Formulas and Tables. Instrument Society of America, Research Triangle Park, North Carolina.
20.Lee TH (1998) The Design of CMOS Radio-Frequency Integrated Circuits. Cambridge University Press, Cambridge, United Kingdom
21.Sullivan CR (1999) Optimal choice for number of strands in a litz-wire transformer winding. IEEE Transactions on Power Electronics 14:283–291
22.Tang Z, Smith B, Schild JH, and Peckham PH (1995) Data transmission from an implantable biotelemeter by load-shift keying using circuit configuration modulator. IEEE Transactions on Biomedical Engineering 42:524–528
23.Wang G, Liu W, Sivaprakasam M, and Kendir GA (2005) Design and analysis of an adaptive transcutaneous power telemetry for biomedical implants. IEEE Transactions on Circuits and Systems-I 52:2109–2117
7. Challenges in Realizing a Chronic High-Resolution Retinal Prosthesis |
149 |
24.Takahashi M, Watanabe K, Sato F, and Matsuki H (2001) Signal transmission system for high frequency magnetic telemetry for an artificial heart. IEEE transactions on Magnetics 372921–372924
25.Bashirullah R, Liu W, Ji Y, Kendir GA, Sivaprakasam M, Wang G, and Pundi B (2003) A smart bi-directional telemetry unit for retinal prosthetic device. In: Proceedings of International Symposium on Circuits and Systems. pp 5–8