Abdul Basir, Youngdae Cho, Izaz Ali Shah, Shahzeb Hayat, Sana Ullah, Muhammad Zada, Syed Ahson Ali Shah, Hyoungsuk Yoo
Editor’s Note
This article for the “Bioelectromagnetics†column takes a deep dive into the field of wireless implants and ingestibles, with a focus on antenna design. The review discusses recent advances in implantable/ingestible antenna design, provides a protocol for the successful design of such antennas, and concludes with an overview of the wide range of real-world applications that may benefit from this research area.
This column welcomes articles on biomedical applications of electromagnetics, antennas, and propagation in terms of research, education, outreach, and more. If you are interested in contributing, please e-mail me at kiourti.1@osu.edu.
Biomedical implantable technologies are life-saving modalities for millions of people globally because of their abilities of wireless remote monitoring, regulating the abnormal functions of internal organs, and early detection of cognitive disorders. Enabling these devices with wireless functionalities, implantable antennas are their crucial front-end components. Detailed overviews of the implantable and ingestible antennas, their types, miniaturization techniques, measurement phantoms, biocompatibility issues, and materials are available in the literature. This article comprehensively reviews the design processes, design techniques and methods, types of antennas, electromagnetic (EM) simulators, and radio-frequency (RF) bands used for implantable and ingestible antennas. We briefly discuss the latest advancements in this field and extend their scope beyond conventional implantable applications. Their related issues and challenges are highlighted, and the performance enhancement techniques have been discussed in detail. All the scoped implantable applications have been covered in this review. A standard protocol has been devised to provide a simple and efficient road map for the design and realization of the implantable and ingestible antenna for future RF engineers and researchers. This protocol minimizes the errors in simulations and measurements by enhancing the agreement between simulated and measured results and simplifies the process of development of implantable and ingestible antennas. It generalizes the process from idea to realization to commercialization and provides an easy road map for the industry.
The enormous developments in biomedical engineering and wireless technologies have advanced our knowledge of human body electrical systems and related phenomena. These significant trends have not only revolutionized health-care systems but have also improved the lifestyle of patients by shifting in-person hospitalization to remote monitoring using implantable wireless medical devices (IWMDs). Typical health-care applications using IWMDs include retinal prostheses, neural recording and stimulation, intracranial pressure measurement, capsule endoscopy, prevention of strokes through stimulation/early detection, blood sugar/glucose level monitoring, hypopnea syndrome diagnosis, cancer treatment, recovery of hearing loss, and endovascular tracking and smart sensing [1], [2], [3], [4], [5], [6], [7], [8], [9], [10], [11], [12], [13], [14], [15], [16], [17], [18]. The IWMD applications can be divided into four subcategories: monitoring, telemetry, stimulation, and treatment. The process of these applications starts with the collection of data by the IWMDs and the transmission of these data wirelessly to on-body devices or the reception of a wireless command from an outside controller to the IWMD. A remote health-care monitoring architecture through IWMDs is shown in Figure 1. As front-end RF components, implantable and ingestible antennas play a crucial role in controlling the operations of IWMDs owing to their inherited wireless performances. Therefore, the performance of these IWMDs depends solely on the acute performance of the implantable antennas. If the implantable antenna fails to perform, the IWMD is rendered useless, thereby risking the lives of patients.
Figure 1. A detailed overview of the implantable and ingestible device applications for remote monitoring, architectures of the IWMDs, and information flow to and from them [1], [2], [3], [4], [5], [6], [7], [8], [9], [10], [11], [12], [13], [14], [15], [16], [17], [18]. MC: microcoil.
The design of implantable antennas starts with the selection of the frequency band. The selection of an appropriate band for real-time telemetry must consider many factors, such as the type of application, size of the implant, implantation depth, implantation site, transmission rate, range of telemetry, patient safety, and the region in which it can be used [19]. Unlicensed bands, such as the medical device radio band (MedRadio: 401–406, 413–419, 426–432, 438–444, 451–457 MHz), Medical Implant Communication System band (MICS: 402–405 MHz), or Industry, Science, Medicine band (ISM band: 433.1–434.8, 868–868.6, 902.8–928, 2,400–2,438.5 MHz), are used for IWMDs [20]. The lower frequency band, that is, the MICS band, is preferred because of lower transmission losses, longer ranges, and ensured tissue safety; however, the bandwidths are narrow and data rates are low at lower bands [21]. Unlike lower frequency bands, higher frequency bands are favorable for high data rates; however, they lead to high transmission losses, a high specific absorption rate (SAR), and shortening of the transmission range. Moreover, miniaturized implantable antennas can be achieved at higher frequencies [22]. Thus, the selection of an appropriate frequency band is crucial and subjective to telemetric needs and types of applications. Deep-body implants use moderate frequency bands to achieve a high data rate with miniaturized antenna geometry and an acceptable transmission range.
Over the past decade, implantable antennas have been in the research spotlight, and advancements have been made to achieve improved performance. Accurate simulations and measurement setups have been devised to standardize the design and realization of these antennas [23]. The design of an antenna within a device-like environment minimizes errors between engineering designs and real-world applications [23]. Further, exploring value-aided techniques (that is, circular polarization, flexible substrates, and metasurfaces) for implantable devices has the potential for improved performance and increased transmission range. Moreover, exploring the irregular effects of heterogeneous tissue environments of the human body on different antennas can lead to overcoming the sensitivity factors of implantable antennas [24]. An overview of the above-mentioned techniques, concepts, and developments can familiarize implantable antenna designers with methods for realizing and standardizing state-of-the-art IWMDs. Therefore, this article presents a comprehensive review of the techniques, design processes, and recent advancements in the field of implantable and ingestible antennas. Table 1 presents an overview of the previous reviews on implantable and ingestible antennas [12], [20], [25], [26], [27], [28], [29], [30], [31], [32] and summarizes the trends and design techniques of implantable and ingestible antennas. As a comparison to start-of-the-art review articles, this review presents the recent trends, a standard design protocol, accurate simulation and measurement setups, and a road map for future development.
Table 1. Summary of the reviews of the implantable and ingestible antennas in the literature and a detailed comparison.
In the literature, antenna topologies including monopoles [37], [38], dipoles [19], [39], [40], [41], loops [22], [35], [42], patches, and planar inverted-F antennas [25], [34], [43], [44], [45], [46], [47], [48], [49], [50] have been employed for implantable and ingestible antennas. To fit these antennas inside a small IWMD, the implantable antenna geometries can be categorized into planar and conformal geometries [22], [33], [36], [51]. Miniaturization strategies, such as spiral shapes, meandering lines, stacked patches, the use of high-permittivity substrates and superstrates, shorting pins, and open-ended ground slots are used to achieve small sizes [1], [22], [23], [43], [44], [52], [53], [54]. Planar antennas have achieved ultraminiaturized geometries to fit inside small IWMDs, such as leadless pacemakers and brain devices, in which the device diameter is less than 6 mm [33], [54], [53]. However, with such small dimensions, only the high-frequency band (2.4 GHz) can be achieved, incurring increased losses, high sensitivity of the antenna, and less efficient systems. Moreover, some of these antennas have ground slots, which increase the sensitivity of these antennas to the onboard circuitries owing to coupling with the printed circuit board (PCB) of the device [2], [3], [4], [33], [36], [53], [55]. On the other hand, conformal antennas experience less of a coupling effect and improved gain because of the allowed large size [22], [56]. These antennas are wrapped inside the package of the IWMDs in such a way as to occupy less space inside the device. Conformal antennas are mostly used in endoscopes or leadless pacemakers, where the devices have cylindrical geometries to provide a smooth curvature for antenna wrapping. Based on miniaturization techniques, achieved bands, geometries, and enhancement techniques, some recent works are shown in Figure 2. Figure 2(a) and (b) shows spiral triple-band and wideband antennas presented, designed, and measured in [33] and [4], respectively. The meandered line tri-band antennas are optimized and tested for small devices, such as the leadless pacemakers in [1] and [34], as shown in Figure 2(c) and (d). Conformal antennas with dual-band circular polarization and wideband characteristics are designed for capsule endoscopes in [22] and presented in Figure 2(e). The authors in [35] have designed a metamaterial-loaded dual-band circular polarized conformal antenna [Figure 2(f)] for a capsule endoscope. Figure 2(g) and (h) shows dual-band circular polarized planar antennas for deep-body implantations [36], [21].
Figure 2. Geometries and related results for different types of implantable and ingestible antennas. (a), (b) Spiral multiband and ultrawideband implantable antennas [33], [4]. (c), (d) Meandered line multiband implantable antennas and their S11 values [1], [34]. (e), (f) Conformal implantable antennas for capsule endoscopes and their S11 values [22], [35]. (g), (h) Dual-band circular polarized planar antennas and their axial ratios [36], [21]. Sim.: simulated; Meas.: measured; BW: bandwidth; HFSS: high-frequency structure simulator; AR: axial ratio.
The design of an implantable antenna is completely different from that of free-space antennas because of the lossy and heterogeneous operating environments [4]. The human body consists of multiple layers (bone, muscle, fat, and skin), and their dielectric properties (conductivities and dielectric constants) vary across the layers and across tissues (termed heterogeneous in the biological sciences). Therefore, the operating environment of the IWMD is different and more complex than that of devices operating in free space. Owing to the heterogeneous and lossy behavior of human tissues, the development of these IWMDs is constrained by certain factors. For example, the communications range and data rates are inevitably reduced because of signal loss in the lossy tissue media and reflections between layers. Moreover, the varying permittivity behavior of the tissues causes impedance mismatches and frequency detuning. These issues complicate the design of antennas for these devices. Furthermore, these devices have limited available space for the device components, meaning that all of the components, including the antenna, must be miniaturized. However, miniaturization makes the antenna narrow band, less efficient, and highly sensitive to the surrounding environment. The performance of these narrow-band antennas is severely affected if the implantation site is changed or a conductor, such as a PCB or battery, is placed near them. In either case, the operation frequency of the antenna shifts to either the higher or lower side, causing communication failure. Therefore, it is necessary to standardize the design process of implantable and ingestible antennas.
For the accurate and stable design of implantable and ingestible antennas for real-life applications, three important factors: device architecture, simulations, and measurement setups, have been discussed in the literature. An overview of each of these follows.
Before 2014, all related studies used a biocompatible coating layer for insulation of the antenna to avoid direct contact with the tissue in both simulations and measurements [55]. However, in real-world implantable applications, the antenna must be placed inside a device with a complete device package. Therefore, the implantable antennas were designed prior to the work by Liu et al. [55] and were rarely used within real devices. In [55], the authors designed a capsule endoscope and an antenna inside it. This study presented interesting effects of the packing on antenna performance. Following this work, in 2016, two device architectures were designed by Gani and Yoo [61] for skin implants. The geometrical parameters of the antenna were modified for the effects of the device. However, neither study measured the antenna within the device; therefore, the simulated and measured results were not matched. To completely analyze the packaging effects on the antenna, Das and Yoo [62] designed a wideband endoscopic antenna, comprehensively analyzed the antenna performance inside a realistic device, and set the trend to follow for future engineers and researchers of implantable antennas.
Various devices have been designed based on commercially available implantable devices for application-oriented implantable antenna design and measurements. For example, Das and Yoo [62] designed and fabricated a capsule-type device with a diameter and length of 10.25 mm and 20.5 mm, respectively, for leadless pacemaker applications. In [3], Shah and Yoo designed two device architectures for scalp-implantable devices with small volumes of 344 mm3 and 406 mm3. Following this trend, the research group led by Prof. Yoo from South Korea and other research groups developed multiple devices, including leadless pacemakers [59], [63], capsule and flat-type devices for deep-body implants [1], [4], [38], [57], [58], [60], [51], [62], and capsule endoscopes [4], [22], [57], [58], [64], [65]. Figure 3 presents a detailed overview of the devices used for simulations and measurements of implantable and ingestible antennas.
Figure 3. The geometries of the full-package devices for the simulations and measurements of the implantable and ingestible antennas. (a) Device for brain-, heart-, and stomach-implantable antennas [21]. (b) Antenna inside a flat-type device for deep implantations [57]. (c), (d), (e) Antennas inside the capsule endoscope [22], [57], [58]. (f), (g) Antennas with leadless and conventional pacemaker systems [59], [60].
Simulations in scientific research speed up the process of device development and reduce time and costs. The studies in literature used finite-element method (FEM) and frequency-difference time domain (FDTD) simulation tools, such as high-frequency structure simulator (HFSS), CST, Xfdtd, and Sim4Life, which are licensed simulators with certain advantages and disadvantages. For example, frequency-domain simulators, such as HFSS, provide fast and efficient optimization of the antenna in single-layered homogeneous phantoms or simple two-to-three-layered phantoms. The simulation time in HFSS depends on the number of layers in the phantom, range of analyzed frequency, size of the phantom, and geometry of the device [33], [36], [49], [51], [54], [53], [55], [61], [62], [66], [64], [65]. On other hand, time-domain-based simulation tools are highly accurate but take a long time. Therefore, a standard method must be followed to reduce the optimization time, obtain precise results, and use fewer computational resources. For a speedy design process and easy follow-up method, we devised a customized standard protocol for developing implantable antennas, as shown in Figure 4. This protocol leads to an accurate and simple realization of implantable antennas, which starts with the aim of the targeted application. Along with the size constraints, the frequency band is selected based on safety and data rate requirements. Then, the architecture of the device is to be modeled to mimic the device-like environment and design the antenna inside the package. Further, the selection of the type and size of homogeneous phantom depends on the application type and implantation depth. Faisal et al. [67], Yousaf et al. [65], and Malik et al. [20] have extensively reviewed the dimensions and types of homogeneous phantoms for implantable antennas. These dimensions range from 25 mm × 25 mm × 25 mm [53] to 200 mm × 200 mm × 200 mm [22]. For deep-body antennas, the electrical properties of muscles (conductivity ${\sigma}$ and permittivity ${\varepsilon}_{r}$) have been used, and the skin properties have been used for the phantoms of skin-implantable antennas. A three-layered model of muscles, fat, and skin is used in [4] and [60]. The frequency-dependent properties of different tissues are defined and tabularized by Gabriel in [68] and [69] and stored on the IT’IS website for the frequency range of 10 Hz–100 GHz [70]. These properties can be assigned to the tissue phantoms in the simulation environments at either a point frequency or range [71], [72].
Figure 4. A road-map protocol for the design and realization of implantable and ingestible antennas. ASTM: American Society for Testing and Materials; ICNIRP: International Commission on Non-Ionizing Radiation Protection.
Once the implantable and ingestible antenna is fully optimized in the frequency-domain simulator, the next step is to verify its performance in FDTD-based heterogeneous environments (Sim4Life by Zurich Med Tech, CST Microwave Studio by CST Studio Suite, and Xfdtd by Remcom) [41], [73], [74]. These FDTD-based commercial simulators have realistic human models, such as Duke, Ella, and Glenn in Sim4Life, GUSTAV in CST, and a heterogeneous human model in Xfdtd. As shown in Step 2 of Figure 4, the antenna performance must be verified using a realistic human model in an FDTD-based EM environment [4]. These simulation tools have been used to verify the antenna performance inside different organs and assess the EM safety in terms of SAR calculations [75]. Basir et al. completely verified the performance and performed a SAR analysis of the capsule antenna in the Duke human model in Sim4Life [22]. Zada and Yoo confirmed the simulation results of a triple-band antenna in different organs (head, colon, heart, stomach, etc.) using a realistic heterogeneous male model in Xfdtd [1]. In [53], Yousaf et al. analyzed a multiband conformal endoscopic antenna in the GUSTAV model of CST, verified its S11, and justified human safety in terms of the SAR.
Although antenna-integrated medical devices placed inside the body offer unique opportunities and enable significant advancements in monitoring, diagnosis, and treatment, there are concerns regarding the safety of these devices owing to their proximity to the user’s body. The International Commission on Non-Ionizing Radiation Protection (ICNIRP) and IEEE have specified basic restrictions in terms of the SAR (rate of EM energy absorbed per unit mass of tissue) to prevent tissue heating due to EM radiation from the aforementioned medical devices. The ICNIRP (2010) basic restrictions confine the SAR averaged over 10 g of contiguous tissue to be less than 2 W/kg, while the IEEE C95.1-1999 standard (2019) limits the SAR average over any 1 g of tissue to be less than 1.6 W/kg [76], [77]. Based on the restricted SAR limits, the effective isotropic radiated power (EIRP) of IWMDs must be limited to ensure safety. Implantable antennas with a high EIRP cause health problems as well as interference with nearby radio devices. An implantable antenna operating in the ISM and MedRadio bands must have an EIRP standard of –20 dB and –16 dB, respectively. Similarly, the input power of the telemetry antenna must be limited to prevent tissue damage. Moreover, the external power source should meet these standards when the implantable antenna operates in the receiver mode [56]. A detailed SAR comparison of various implantable and ingestible antennas and their corresponding radiated power limits is presented in Table 2.
Table 2. SAR and Maximum Power Levels for Compliance with Safety Standards of various Implantable and ingestible antennas.
After optimization of the implantable and ingestible antenna inside a full-package device in a homogeneous model and followed by its performance verification in an FDTD-based simulation environment in a heterogeneous human model, the antenna is ready for fabrication and measurement. The measurement process starts by connecting a flexible insulated subminiature-A connector with a suitable length and a diameter of 1 mm. The antenna is encapsulated in a 3D-printed device (capsule or flat type) made of biocompatible material and implanted into the phantom for in vitro measurement. The device must be sealed such that the liquid from the phantom does not leak into the encapsulated antenna [2], [22]. Step 3 of the protocols (Figure 4) shows the in vitro validation method for implantable and ingestible antennas. In the literature, different phantoms are used as measurement setups. These phantoms can be categorized as gel based, such as saline solution [23], semisolid and solid chemical-based phantoms, and dead animal tissue based, such as ground beef, minced meat, and porcine heart placed in a saline-filled American Society for Testing and Materials (ASTM) phantom. The semisolid and solid chemical-based phantoms were comprehensively reviewed in [31], and the standard saline-based and minced meat phantoms are shown in Figure 5. A saline-filled 3D head phantom, which has the dimensions of an average young Chinese male, was used for measurements of head-implanted antennas [3], [23]. A saline-filled ASTM phantom with dimensions comparable to those of the human torso was used for the measurement of capsule endoscopes [2], [22]. Moreover, containers filled with minced pork are used for the measurements of other implantable and ingestible antennas in [22]. Because of their nearly matched dielectric properties with those of human tissues, minced pork phantoms provide accurate and realistic results [78]. Additionally, the permittivity and conductivity can be easily controlled by adding an accurate fat ratio during the grinding process.
Figure 5. Setups for S11 and gain measurements and telemetry of implantable and ingestible antennas. (a) A saline-filled ASTM phantom for measurement of capsule antenna [22]. (b) A saline-filled box for measurement of implantable and ingestible antennas. (c) Meat phantom for skin-implantable antennas [60]. (d) Minced meat phantom for measurement of MIMO ingestible antennas [57]. (e) A 3D-printed head phantom filled with saline solution for measurement of MRI-assisted implantable antenna [73]. (f), (g) Radiation pattern measurement setups for implantable and ingestible antennas [59], [22]. (h) Testbed for checking real-time telemetric capabilities [22]. MIMO: multiple-input, multiple-output; MRI: magnetic resonance imaging; AUT: antenna under test; SDR: software-defined radio; Tx: transmitter; Rx: receiver.
After satisfying the aforementioned steps of the design protocol, the antenna is ready for use in in vivo testing. Very few studies have reported in vivo measurements of the scattering parameters of implantable and ingestible antennas. However, in vivo antennas are always used with the wireless module to perform telemetry as a standalone wireless system. The open surgery time for device implantation is limited, and the parameters of these antennas, such as the radiation pattern and S11, etc., are not required to be measured in-vivo. Instead, the matching and integration of the antenna with the device are required. For this, an approach was devised in [22] and [57]. For example, as shown in Figure 5(h), a software-defined radio (SDR), with a standard impedance of 50 Ω and working as a transmitter, was connected to the ingestible antenna implanted in a saline-filled ASTM phantom. Furthermore, a simple monopole antenna was connected to the same type of SDR, working as a receiver, and the data transmission in real time was tested. The SDR through the implantable and ingestible antenna, placed inside the phantom, successfully communicated and sent real-time data to the SDR connected to the monopole near the ASTM phantom.
Magnetic resonance imaging (MRI) provides quantitative details of the imaging object, and recently antenna designs for MRI have been proposed owing to the strong EM field generation at resonance frequencies [79], [80], [81], [82], [83]. Deep blood vessel imaging and structural information are crucial in fatal diseases, such as vascular plaque and intracranial brain aneurysms; however, the imaging resolution of these vessels is extremely low [83]. Implantable and ingestible antennas used in MRI are termed microcoils (MCs). MCs are recommended when the patient’s blood vessel images have a low resolution and signal-to-noise ratio (SNR). A small catheter-based MC antenna attached to the tip of the catheter wire provides MR images of the blood vessels and surrounding tissues. An MC antenna can be implanted in the target area using a vascular access procedure.
MC antennas are widely used in deep brain tissues and blood vessels. Therefore, MC antennas are designed considering the environmental effects of the deep tissues and blood vessels, such as small diameter, lossy human tissues, flowing blood, blood vessel walls, and catheter lead wire. Because the diameter of the deep vessel is very small, miniaturized and compact MC structures resonating at lower megahertz bands (MRI resonance frequencies) are required for implantation in the target location. Researchers have proposed single-loop and opposite-solenoid designs integrated with catheter wire to capture the luminal images of vessels [84], [85], [86], [87].
However, a few small vessels exist in which the loop and solenoid are not practical and require ultracompact MC antenna designs. Therefore, loopless catheter antennas, such as monopole and dipole antennas, are used to visualize small-diameter vessels [87], [88]. Although these MC designs provide a higher SNR and better visualization, the dipole length and MC placement in the blood vessel remain major issues. The alternative solution and best design for use inside blood vessels is the hollow-shaped MC antenna, which can be perfectly placed in the blood vessel lumen [89], [90], [91], [92].
Another major challenge in MC antenna design is the blood flow inside the vessel, which creates an artifact in the MR image and increases the noise figure. The previously designed MCs, such as single-loop, solenoid, tilted, saddle-shaped, and meandering line MCs, overcame the SNR issues of deep brain tissues and the interior of blood vessels; however, the blood flow blockage due to MC is not resolved in these designs. As the coil moves during intense blood flow, severe image artifacts are often observed in the arterial lumen rather than near the arterial walls [73], [92]. Therefore, an MC design with a hollow shape that is easily wrapped inside the vessel and completely covers the target area of the blood vessel is required for a large field of view, higher SNR, and homogeneous field distribution inside the blood vessel. A detailed overview of the different MCs is presented in Table 3.
Table 3. Qualitative comparison of implantable and ingestible MC antennas.
Owing to the small size and lossy tissue operation environment, the gain and efficiency of the implantable and ingestible antennas are small [21], [93]. The Friis equations show that improvement in gain strengthens the wireless communication link [63]. Therefore, only a few studies have adopted techniques to enhance the gain and efficiency of implantable and ingestible antennas [21], [94], [35]. In [94], the authors have used a deionized water-infilled cavity on the back of a scalp-implantable antenna to increase its gain and efficiency. However, works have extended applications of metamaterials to implantable antennas [21], [35]. Zada et al. [21] and Samanta and Mitra [35] loaded their antennas with double-positive (with positive effective permittivity and permeability) metamaterials to enhance gain and efficiency. Both works achieved at least a 1.5-dB improvement in gain and efficiency.
Because of the use of lower frequencies, where the available bandwidths are narrow, the data rates and spectral efficiency of current IWMDs integrated with single-input-single-output antennas are low. As a solution, multiple-input, multiple-output (MIMO) antennas have been introduced to improve the channel capacity and satisfy the data rate requirements of modern IWMDs. Two-element MIMO antennas in [57], [95], [96], and [97] and four-element MIMOs in [58], [98], [99], [100], and [101] were designed and measured for capsule endoscopic and skull-implantable device applications.
Wireless multitasking, such as full-duplex telemetry, simultaneous data telemetry, and wireless battery charging, requires extra multiplexer circuitry located just after the implantable and ingestible antennas. However, these multiplexers have a complex functionality and add an extra circuitry burden to the devices at the cost of making the device heavy and requiring more power consumption. To simplify the device design and remove the external multiplexer, a circular diplexer antenna with a diameter of 9.6 mm is designed in [64]. The diplexer antenna has two closely placed half-circled patches with two independent ports, one for uplink telemetry and the other for downlink wireless power transfer. Multiplexer antennas can significantly simplify the circuitry and functions of IWMDs and extend their lives by reducing power consumption.
The efficient monitoring of patients with abdominal aortic aneurysm and the sharing of biotelemetry data comprising information about certain physiological indicators are life-saving procedures [12]. Smart stents featuring antenna characteristics are promising candidates for endovascular aneurysm repair (EVAR) and as antennas to transmit telemetric data to the external controller. Shah et al. developed a novel endovascular aortic stent system that possesses promising antenna characteristics and has the potential to be used for biotelemetry in EVAR applications [46]. Moreover, biotelemetric-enabled smart stents integrated with various sensors, including blood-flow sensors, have been presented in [12], [102], [103], and [104].
This article presents an overview of implantable and ingestible antennas for device integration as well as their related issues. Common geometries and types of antennas designed for implantable devices, simulation environments, measurement techniques, performance enhancement methods, and novel applications of implantable and ingestible antennas existing in the literature are summarized. Furthermore, all of the issues related to implantable and ingestible antennas are comprehensively discussed, and important works are overviewed. The geometries of the antennas, device types, simulation and measurement setups, and techniques are categorized based on their types and applications. Each classification is discussed in terms of practicality, feasibility, and its advantages and disadvantages. A detailed study on human safety under EM exposure for implantable and ingestible antennas is presented. Furthermore, a protocol is devised as an easy follow-up standardized method for new researchers in this field and related industries. According to this new protocol, the first step is the selection of an appropriate frequency band, designing the full package device for the antenna, and starting the optimization of the antenna inside the device. This step minimizes the gap between the design and prototyping for real-world applications. This standard protocol is a complete guideline for the design and realization of implantable and ingestible antennas as an explicit road map from thinking to commercialization. Moreover, performance enhancement techniques, such as the gain enhancement technique, data rate enhancement (using MIMO), and simultaneous multitasking (multiplexer antennas), are discussed.
This work was supported by a National Research Foundation of Korea grant funded by the Korean Government Ministry of Science and ICT (MSIT) (Grant 2022R1A2C2003726) and the Institute of Information and Communications Technology Planning and Evaluation funded by MIST, under Grant 2022-0-00310. The corresponding author is Hyoungsuk Yoo. Abdul Basir and Youngdae Cho are co-first authors and contribute equally.
Abdul Basir (engrobasir@gmail.com) is a postdoctoral researcher in the Department of Electronic Engineering, Hanyang University, Seoul 04763, South Korea. His research interests include implantable antennas and systems, biomedical circuits, metamaterials, dielectric resonator antennas, reconfigurable antennas, long-range wireless power transfer, and stretchable electronics. He is a Member of IEEE.
Youngdae Cho (chb1046@hanyang.ac.kr) is a postdoctoral researcher in the Department of Electronic Engineering, Hanyang University, Seoul 04763, South Korea. His research interests include implantable antennas and devices, wireless power transfer, magnetic resonance imaging, and radio-frequency heating and safety. He is a Member of IEEE.
Izaz Ali Shah (izazaliuet@gmail.com) is a postdoctoral researcher in the Department of Electronic Engineering with the Applied Bioelectronics Laboratory, Hanyang University, Seoul 04763, South Korea. His research interests include implantable antennas and devices, wireless power transfer, implant safety in high magnetic field systems, and frequency-selective surfaces. He is a Graduate Student Member of IEEE.
Shahzeb Hayat (shahzebuet@gmail.com) is a Ph.D. candidate in the Department of Electronic Engineering at the Applied Bioelectronics Laboratory, Hanyang University, Seoul 04763, South Korea. His research interests include implantable antennas and systems, reconfigurable antennas, magnetic resonance imaging (MRI) and radio-frequency coils, and intravascular catheter tracking under MRI. He is a Graduate Student Member of IEEE.
Sana Ullah (sanaullah.9457841@gmail.come) is a postdoctoral researcher in the Department of Electronic Engineering, Hanyang University, Seoul 04763, South Korea. His research interests include 60-GHz antennas, implantable antennas, radio-frequency (RF) coils for magnetic resonance imaging (MRI), wireless MRI, and flexible RF coils. He is a Graduate Student Member of IEEE.
Muhammad Zada (muhammadzada21@gmail.com) is a postdoctoral fellow at Hanyang University, Seoul 04763, South Korea. His research interests include implantable antennas, intra-oral tongue drive systems, wireless power transfer, millimeter-wave antennas, wearable sensors, smart textiles, and microwave cancer detection. He is a Graduate Student Member of IEEE.
Syed Ahson Ali Shah (sahsonas@gmail.com) is a postdoctoral researcher at Tesla Laboratory at Gwangju Institute of Technology, Gwangju 61005, South Korea. His research interests include implantable antennas and systems, wireless power transfer to biomedical devices, sensor-integrated telemetric stents, and metamaterial-based antenna systems. He is a Graduate Student Member of IEEE.
Hyoungsuk Yoo (hsyoo@hanyang.ac.kr) is a full professor in the Department of Biomedical Engineering and the Department of Electronic Engineering, Hanyang University, Seoul 04763, South Korea. He is the CEO of E2MR. His research interests include electromagnetic theory, metamaterials, antennas, and magnetic resonance imaging in high-magnetic field systems. He is a Senior Member of IEEE.
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Digital Object Identifier 10.1109/MAP.2023.3301398
