A Single-Connector Stent Antenna for Intravascular Monitoring Applications

Recently, smart stents have been developed by integrating various sensors with intravascular stents for detecting vascular restenosis or monitoring intravascular biomedical conditions such as blood pressure or blood flow velocity. The information on biomedical signals is then transmitted to external monitoring systems via wireless communications. Due to the limited volumes of blood vessels and limited influence of blood flow, antennas with good radiation performance are required for intravascular applications. In this paper, we propose a stent antenna composed of multiple rings containing crowns and struts, where each ring is connected with one connector. Unlike a conventional stent, wherein each ring is connected with several connectors, the single connector prevents the random distribution of electrical current and thus achieves good radiation performance. The implantable stent antenna is designed for the frequency range of 2 to 3 GHz for minimum penetration loss in the human body and tissues. Mechanical FEM simulations were conducted to ensure that the mechanical deformation was within specific limits during balloon expansions. A prototype was fabricated with laser cutting techniques and its radiation performance experimentally characterized. It was demonstrated that the fabricated stent antenna had an omnidirectional radiation pattern for arbitrary receiving angles, a gain of 1.38 dBi, and a radiation efficiency of 74.5% at a resonant frequency of 2.07 GHz. The main contribution of this work was the manipulation of the current distributions of the stent for good EM radiation performances which needed to be further examined while inserted inside human bodies. These research results should contribute to the further development of implantable wireless communications and intravascular monitoring of biomedical signals such as blood pressure and blood flow velocity.


Introduction
In past decades, intravascular stents have been widely used in the medical treatment of vessel obstructions for patients with vascular diseases such as stroke, heart attack, and aneurysm [1,2]. Stents are essentially wire meshes inserted into narrow blood vessels via a catheter and medical surgery. The narrow parts of the blood vessels are expanded by balloons to restore regular blood flow. However, the long-term usage of stents inside the blood vessels leads to vascular restenosis (i.e., vessel re-narrowing) due to the growth of vascular smooth muscle cells [3]. The re-narrowing of blood vessels hinders the blood flow and can cause severe problems. These issues can be alleviated by eluting drugs [4,5], manipulating shear stresses on the vessel walls [6], or heating the re-narrowing regions to suppress the growth of vascular smooth muscle cells [7,8].
For early diagnosis of restenosis and real-time monitoring of intravascular blood conditions, smart stents or intelligent stents have been developed by integrating versatile sensors and wireless In general, intravascular stents are small in size, ranging in length from 12 to 38 mm and in diameter from 2 to 10 mm. Thus, it is necessary to miniaturize implantable electronic devices and communication systems. Various wireless power/data communications approaches for smart stents have been investigated, including passive LC-resonant antennas [7][8][9]12,17,[24][25][26], miniaturized packaged antennas [20], RFIC [11], and active stent antennas [10,14,15,19,23]. LC-resonant antennas are composed of metallic stents acting as inductors and additional capacitors forming LC resonant circuits. Via strong magnetic field couplings with external loop antennas and the phase dip technique, the intravascular information can be obtained through the variations of the resonant frequencies.
However, one disadvantage of LC resonant antennas is that the active regions of the transmitting loop antennas should be overlapped with the receiving inductive stents for strong magnetic-field couplings [17]. In some circumstances, it is difficult to locate coronary artery stents by the commonly used CT angiography [27] or MRI [28]. However, the whole stent can be considered as an antenna for receiving electromagnetic signals and power from transmitting antennas at arbitrary angles [23]. For stent antennas, the requirements that need to be met are long-term usage, biocompatibility, electrical conduction, and radial expansion capability, as well as structural integrity after expansion to prevent the influence of blood flow.
In this paper, we propose a stent antenna for intravascular monitoring and implantable wireless applications. The stent antenna was expected to have better radiation performance than that of LC resonant antennas at low frequencies, such as 900 MHz [19]. Modern stents are composed of open cells, each unit is composed of crowns and struts, and each ring is connected by multiple connectors [29]. However, this design would cause random current distribution and reduce radiation performance. The goal of this work was to design a stent antenna with good radiation performance by using a single connector between each unit of the stent. The single-connector stent antenna eliminates the complicated and meandering geometry of a typical stent, which would cancel out the induced currents and be unable to excite the desired electromagnetic waves. In addition, a stent with single connectors should maintain a stable mechanical structure and not exceed fracture limits during In general, intravascular stents are small in size, ranging in length from 12 to 38 mm and in diameter from 2 to 10 mm. Thus, it is necessary to miniaturize implantable electronic devices and communication systems. Various wireless power/data communications approaches for smart stents have been investigated, including passive LC-resonant antennas [7][8][9]12,17,[24][25][26], miniaturized packaged antennas [20], RFIC [11], and active stent antennas [10,14,15,19,23]. LC-resonant antennas are composed of metallic stents acting as inductors and additional capacitors forming LC resonant circuits. Via strong magnetic field couplings with external loop antennas and the phase dip technique, the intravascular information can be obtained through the variations of the resonant frequencies.
However, one disadvantage of LC resonant antennas is that the active regions of the transmitting loop antennas should be overlapped with the receiving inductive stents for strong magnetic-field couplings [17]. In some circumstances, it is difficult to locate coronary artery stents by the commonly used CT angiography [27] or MRI [28]. However, the whole stent can be considered as an antenna for receiving electromagnetic signals and power from transmitting antennas at arbitrary angles [23]. For stent antennas, the requirements that need to be met are long-term usage, biocompatibility, electrical conduction, and radial expansion capability, as well as structural integrity after expansion to prevent the influence of blood flow.
In this paper, we propose a stent antenna for intravascular monitoring and implantable wireless applications. The stent antenna was expected to have better radiation performance than that of LC resonant antennas at low frequencies, such as 900 MHz [19]. Modern stents are composed of open cells, each unit is composed of crowns and struts, and each ring is connected by multiple connectors [29]. However, this design would cause random current distribution and reduce radiation performance. The goal of this work was to design a stent antenna with good radiation performance by using a single connector between each unit of the stent. The single-connector stent antenna eliminates the complicated and meandering geometry of a typical stent, which would cancel out the induced currents and be unable to excite the desired electromagnetic waves. In addition, a stent with single connectors should maintain a stable mechanical structure and not exceed fracture limits during balloon expansion [1]. The single-connector stent antenna could be applied to other vessels and not limited to coronary arteries. This paper is organized as follows. In the next section, the design of the proposed stent antenna and its operation frequency are presented. Then, the EM and FEM simulations determining the electrical properties and mechanical FEM simulations for verifying the mechanical properties are described. Subsequently, the laser-cutting fabrication process and balloon expansions of prototypes are disclosed. Finally, the radiation characteristics of the fabricated stent antenna are measured in an anechoic chamber and discussed. Important results are summarized at the end of the paper. Figure 2a presents the 2D unfolded planar geometry of a multi-connector stent. In general, multi-connector stents are composed of several rings connected by connectors with cylindrical shapes for providing sufficient radial force to open narrowed blood vessels [5,24]. Each ring consists of crowns and struts, and the crowns are circular so that they can expand and achieve large deformation under balloon expansions. Most commercially available multi-connector stents have two to three connectors between adjacent rings to achieve the desired structural stability [30,31]. However, these multiple connectors can lead to random current distributions and are not suitable for antenna applications. In this paper, we propose a single-connector stent antenna having adjacent rings connected by a single connector, as shown in Figure 2b. A single-connector stent antenna can allow current distribution with an approximate half-wavelength resonant mode. The current distributions will be discussed in the next section. Figure 3a,b presents the 3D geometries of a multi-connector and the proposed single-connector stent antennas where undesired connectors are not connected. balloon expansion [1]. The single-connector stent antenna could be applied to other vessels and not limited to coronary arteries. This paper is organized as follows. In the next section, the design of the proposed stent antenna and its operation frequency are presented. Then, the EM and FEM simulations determining the electrical properties and mechanical FEM simulations for verifying the mechanical properties are described. Subsequently, the laser-cutting fabrication process and balloon expansions of prototypes are disclosed. Finally, the radiation characteristics of the fabricated stent antenna are measured in an anechoic chamber and discussed. Important results are summarized at the end of the paper. Figure 2a presents the 2D unfolded planar geometry of a multi-connector stent. In general, multiconnector stents are composed of several rings connected by connectors with cylindrical shapes for providing sufficient radial force to open narrowed blood vessels [5,24]. Each ring consists of crowns and struts, and the crowns are circular so that they can expand and achieve large deformation under balloon expansions. Most commercially available multi-connector stents have two to three connectors between adjacent rings to achieve the desired structural stability [30,31]. However, these multiple connectors can lead to random current distributions and are not suitable for antenna applications. In this paper, we propose a single-connector stent antenna having adjacent rings connected by a single connector, as shown in Figure 2b. A single-connector stent antenna can allow current distribution with an approximate half-wavelength resonant mode. The current distributions will be discussed in the next section. Figure 3a,b presents the 3D geometries of a multi-connector and the proposed singleconnector stent antennas where undesired connectors are not connected.    For implantable antennas, the resonant frequency of the stent antenna depends on the total length of the stent and its geometry. In fact, in the human body, electromagnetic waves (EM) have different penetration attenuations and variable power loss boundaries at different frequencies [23,32]. For low-frequency electromagnetic coupling communications, the optimal operation frequency is in the MHz range for low tissue absorption loss where the biological tissues are modeled as dispersive medium and results in large sizes of implanted antennas [23]. For high-frequency wireless communications, the human body is modeled as a low-loss multi-layer dielectric medium and the optimal frequency is in the range from 2 to 3 GHz for the minimum tissue absorptions and good orientation tolerances of implanted antennas [32]. Therefore, the single-connector stent antenna was designed to have a resonant frequency in the range of 2 to 3 GHz that the total length of the surface current flow path shown in Figure

Designs
where λ was the wavelength of the first resonance. The strut length, crown length, and connector length are the lengths of the strut, crown, and connector shown in Figure 2. Furthermore, the length should be in the range of 10 to 20 mm, and the diameter, in the range of 2 to 10 mm due to the sizes of vessels [7,8]. In addition, this frequency range falls in the range of the ISM band, and a high operation frequency is beneficial for implantable antennas [23]. The final dimensions of the singleconnector stent antenna were a length of 18 mm, 6 crowns, 9 rings, and a final diameter of 2 mm when inserted into blood vessels, as shown in Figure 4 and Table 1. For comparison, a multiconnector stent antenna and an old helix-shaped stent antenna of the same size were also investigated. The helical stent antenna had 9 turns within the same total length of 18 mm.  For implantable antennas, the resonant frequency of the stent antenna depends on the total length of the stent and its geometry. In fact, in the human body, electromagnetic waves (EM) have different penetration attenuations and variable power loss boundaries at different frequencies [23,32]. For low-frequency electromagnetic coupling communications, the optimal operation frequency is in the MHz range for low tissue absorption loss where the biological tissues are modeled as dispersive medium and results in large sizes of implanted antennas [23]. For high-frequency wireless communications, the human body is modeled as a low-loss multi-layer dielectric medium and the optimal frequency is in the range from 2 to 3 GHz for the minimum tissue absorptions and good orientation tolerances of implanted antennas [32]. Therefore, the single-connector stent antenna was designed to have a resonant frequency in the range of 2 to 3 GHz that the total length of the surface current flow path shown in Figure 2b was close to a half wavelength expressed as where λ was the wavelength of the first resonance. The strut length, crown length, and connector length are the lengths of the strut, crown, and connector shown in Figure 2. Furthermore, the length should be in the range of 10 to 20 mm, and the diameter, in the range of 2 to 10 mm due to the sizes of vessels [7,8]. In addition, this frequency range falls in the range of the ISM band, and a high operation frequency is beneficial for implantable antennas [23]. The final dimensions of the single-connector stent antenna were a length of 18 mm, 6 crowns, 9 rings, and a final diameter of 2 mm when inserted into blood vessels, as shown in Figure 4 and Table 1. For comparison, a multi-connector stent antenna and an old helix-shaped stent antenna of the same size were also investigated. The helical stent antenna had 9 turns within the same total length of 18 mm.  In this research, we employed the commonly used L-605 Cobalt-Chromium (Co-Cr) alloy as the material for all the stents because Co-Cr alloy has a smaller minimum thickness than conventional stainless steel [1]. As a ductile material, Co-Cr alloy has a Young's modulus of 203 GPa, a Poisson's ratio of 0.3, yield stress of 590 MPa, and ultimate strength of 1689 MPa.

EM FEM Simulations
To investigate the EM properties of stent antennas, full-wave EM simulations were conducted to study the surface currents and radiation patterns with commercial software, CST Microwave Studio. Figure 5 shows the surface current distributions at the resonant frequency of the helical, multiconnector, and single-connector stent antennas. Most stent antennas had an issue of small ground planes due to limited small volumes of blood vessels [10,23]. Similarly, we used SMA connectors with finite ground planes for power extractions. It could be improved by combining two stent antennas to form a dipole-like antenna [23]. The multi-connector stent antenna had three connectors, and the random current flowed throughout the entire stent. In the single-connector stent antenna, there was a major current flow path between each ring, resulting in an approximate half-wavelength resonant mode similar to that of the helical stent antenna with a relatively small ground plane. Therefore, the single-connector stent antenna and the helical stent antenna had similar resonant frequencies, as can be observed from the simulated reflection coefficients in Figure 6. Since the multi-connector stent antenna had random current distribution, it had a higher resonant frequency with the same dimensions. This result supported that a single-connector stent antenna is suitable for stent antennas. In this research, we employed the commonly used L-605 Cobalt-Chromium (Co-Cr) alloy as the material for all the stents because Co-Cr alloy has a smaller minimum thickness than conventional stainless steel [1]. As a ductile material, Co-Cr alloy has a Young's modulus of 203 GPa, a Poisson's ratio of 0.3, yield stress of 590 MPa, and ultimate strength of 1689 MPa.

EM FEM Simulations
To investigate the EM properties of stent antennas, full-wave EM simulations were conducted to study the surface currents and radiation patterns with commercial software, CST Microwave Studio. Figure 5 shows the surface current distributions at the resonant frequency of the helical, multi-connector, and single-connector stent antennas. Most stent antennas had an issue of small ground planes due to limited small volumes of blood vessels [10,23]. Similarly, we used SMA connectors with finite ground planes for power extractions. It could be improved by combining two stent antennas to form a dipole-like antenna [23]. The multi-connector stent antenna had three connectors, and the random current flowed throughout the entire stent. In the single-connector stent antenna, there was a major current flow path between each ring, resulting in an approximate half-wavelength resonant mode similar to that of the helical stent antenna with a relatively small ground plane. Therefore, the single-connector stent antenna and the helical stent antenna had similar resonant frequencies, as can be observed from the simulated reflection coefficients in Figure 6. Since the multi-connector stent antenna had random current distribution, it had a higher resonant frequency with the same dimensions. This result supported that a single-connector stent antenna is suitable for stent antennas.

Mechanical FEM Simulations
To verify that the mechanical properties of the stent antennas fulfilled the requirements, mechanical finite element simulations were conducted in commercial software, ABAQUS (Dassault System Simulia Corp., Velizy Villacoublay, France). In accordance with the practical balloon expansion processes of stents, a cylindrical stent with an initial diameter of 1.5 mm was expanded to a diameter of 2.2 mm to simulate balloon expansion, as shown in Figure 7 [33]. Then the oversized stent was compressed to the desired diameter of 2 mm to simulate the radial force of the blood vessel walls. The simulated Von-Mises stress distributions of the multi-connector and single-connector stents are shown in Figure 8. In general, coronary stents use ductile materials that can achieve large deformation during balloon expansion [34]. They obey the Von-Mises stress criterion, expressed as

Mechanical FEM Simulations
To verify that the mechanical properties of the stent antennas fulfilled the requirements, mechanical finite element simulations were conducted in commercial software, ABAQUS (Dassault System Simulia Corp., Velizy Villacoublay, France). In accordance with the practical balloon expansion processes of stents, a cylindrical stent with an initial diameter of 1.5 mm was expanded to a diameter of 2.2 mm to simulate balloon expansion, as shown in Figure 7 [33]. Then the oversized stent was compressed to the desired diameter of 2 mm to simulate the radial force of the blood vessel walls. The simulated Von-Mises stress distributions of the multi-connector and single-connector stents are shown in Figure 8. In general, coronary stents use ductile materials that can achieve large deformation during balloon expansion [34]. They obey the Von-Mises stress criterion, expressed as

Mechanical FEM Simulations
To verify that the mechanical properties of the stent antennas fulfilled the requirements, mechanical finite element simulations were conducted in commercial software, ABAQUS (Dassault System Simulia Corp., Velizy Villacoublay, France). In accordance with the practical balloon expansion processes of stents, a cylindrical stent with an initial diameter of 1.5 mm was expanded to a diameter of 2.2 mm to simulate balloon expansion, as shown in Figure 7 [33]. Then the oversized stent was compressed to the desired diameter of 2 mm to simulate the radial force of the blood vessel walls. The simulated Von-Mises stress distributions of the multi-connector and single-connector stents are shown in Figure 8. In general, coronary stents use ductile materials that can achieve large deformation during balloon expansion [34]. They obey the Von-Mises stress criterion, expressed as where σ 1 , σ 2 , and σ 3 are principle stresses. The Von-Mises stress criterion indicates that ductile stents may fracture if the maximum Von-Mises stress exceeds the ultimate strength. The maximum Von-Mises stresses of the multi-connector and single connector stents in this research were lower than the ultimate strength of 1689 MPa of Co-Cr alloy. In other words, both stents had suitable mechanical properties for balloon expansions. In addition, the effect of removing the connector was described in Figure 9.
If the connector was completely removed, it caused a longer longitudinal deformation, L2, than the original length of the connector, L1, resulting in undesired longitudinal elongations. This could be modified via partially removing the connector shown in Figure 9b that the deformed length, L3, was shorter than L2. Therefore, the trade-off design of the single connector was to sacrifice the longitudinal elongation and gain a better radiation efficiency. In other words, both stents had suitable mechanical properties for balloon expansions. In addition, the effect of removing the connector was described in Figure 9. If the connector was completely removed, it caused a longer longitudinal deformation, L2, than the original length of the connector, L1, resulting in undesired longitudinal elongations. This could be modified via partially removing the connector shown in Figure 9b that the deformed length, L3, was shorter than L2. Therefore, the trade-off design of the single connector was to sacrifice the longitudinal elongation and gain a better radiation efficiency.    In other words, both stents had suitable mechanical properties for balloon expansions. In addition, the effect of removing the connector was described in Figure 9. If the connector was completely removed, it caused a longer longitudinal deformation, L2, than the original length of the connector, L1, resulting in undesired longitudinal elongations. This could be modified via partially removing the connector shown in Figure 9b that the deformed length, L3, was shorter than L2. Therefore, the trade-off design of the single connector was to sacrifice the longitudinal elongation and gain a better radiation efficiency.    In other words, both stents had suitable mechanical properties for balloon expansions. In addition, the effect of removing the connector was described in Figure 9. If the connector was completely removed, it caused a longer longitudinal deformation, L2, than the original length of the connector, L1, resulting in undesired longitudinal elongations. This could be modified via partially removing the connector shown in Figure 9b that the deformed length, L3, was shorter than L2. Therefore, the trade-off design of the single connector was to sacrifice the longitudinal elongation and gain a better radiation efficiency.

Fabrication
Two prototypes of single-connector and multi-connector stents were fabricated via laser cutting of cylindrical Co-Cr alloy hypotubes. To pattern 2D unfolded geometries on the surfaces of the hypotubes, a 30W Nd-YAG pulse type fiber optical laser (Rofin) and an XY movement platform (Aerotech position) with precise position controls were employed. The details of the fabrication processes are reported in [1]. Figure 10 presents photographs of the fabricated multi-connector and single-connector stent antennas connected with SMA connectors for measurements. High-resolution photographs of the fabricated multi-connector and single-connector stent antennas viewed under an optical microscope are shown in Figure 11. In the single-connector stent antenna, each ring was connected via a single connector, and undesired connectors were not connected, as highlighted with a red box.

Fabrication
Two prototypes of single-connector and multi-connector stents were fabricated via laser cutting of cylindrical Co-Cr alloy hypotubes. To pattern 2D unfolded geometries on the surfaces of the hypotubes, a 30W Nd-YAG pulse type fiber optical laser (Rofin) and an XY movement platform (Aerotech position) with precise position controls were employed. The details of the fabrication processes are reported in [1]. Figure 10 presents photographs of the fabricated multi-connector and single-connector stent antennas connected with SMA connectors for measurements. High-resolution photographs of the fabricated multi-connector and single-connector stent antennas viewed under an optical microscope are shown in Figure 11. In the single-connector stent antenna, each ring was connected via a single connector, and undesired connectors were not connected, as highlighted with a red box.

Reflection Coefficients
The reflection coefficients of the fabricated single-connector and multi-connector stent antennas were measured with a calibrated vector network analyzer (Agilent E5071C). To reduce the effect of the feed line, an RF choke (Jin Hua Electronic Co., LTD, model: 2030000001866, Taipei, Taiwan) was utilized during the measurements. Figure 12 shows the simulated and measured reflection coefficients of both stent antennas. The resonant frequency of the single-connector stent antenna was lower than that of the multi-connector stent antenna of the same length. The simulation results matched the measured results, and the slight discrepancy was due to fabrication errors and tolerances.

Fabrication
Two prototypes of single-connector and multi-connector stents were fabricated via laser cutting of cylindrical Co-Cr alloy hypotubes. To pattern 2D unfolded geometries on the surfaces of the hypotubes, a 30W Nd-YAG pulse type fiber optical laser (Rofin) and an XY movement platform (Aerotech position) with precise position controls were employed. The details of the fabrication processes are reported in [1]. Figure 10 presents photographs of the fabricated multi-connector and single-connector stent antennas connected with SMA connectors for measurements. High-resolution photographs of the fabricated multi-connector and single-connector stent antennas viewed under an optical microscope are shown in Figure 11. In the single-connector stent antenna, each ring was connected via a single connector, and undesired connectors were not connected, as highlighted with a red box.

Reflection Coefficients
The reflection coefficients of the fabricated single-connector and multi-connector stent antennas were measured with a calibrated vector network analyzer (Agilent E5071C). To reduce the effect of the feed line, an RF choke (Jin Hua Electronic Co., LTD, model: 2030000001866, Taipei, Taiwan) was utilized during the measurements. Figure 12 shows the simulated and measured reflection coefficients of both stent antennas. The resonant frequency of the single-connector stent antenna was lower than that of the multi-connector stent antenna of the same length. The simulation results matched the measured results, and the slight discrepancy was due to fabrication errors and tolerances.

Reflection Coefficients
The reflection coefficients of the fabricated single-connector and multi-connector stent antennas were measured with a calibrated vector network analyzer (Agilent E5071C). To reduce the effect of the feed line, an RF choke (Jin Hua Electronic Co., LTD, model: 2030000001866, Taipei, Taiwan) was utilized during the measurements. Figure 12 shows the simulated and measured reflection coefficients of both stent antennas. The resonant frequency of the single-connector stent antenna was lower than that of the multi-connector stent antenna of the same length. The simulation results matched the measured results, and the slight discrepancy was due to fabrication errors and tolerances.

Far-field Radiation Characteristics
The 3D far-field radiation patterns were measured in the anechoic chamber at the Joint Laboratory for Microwave and Millimeter Wave Antenna Measurements at National Taiwan University of Science and Technology. Figure 13 shows the measurement setup. The stent antenna was mounted on a rotating stage and acted as a receiving antenna, and the probe antenna, which acted as a transmitting antenna, moved in the azimuth and elevation planes. The measured gain radiation patterns are shown in Figures 14 and 15. Note that the measured radiation patterns at 180° were inaccurate due to the measurement setup where the transmitting antenna couldn't move to the position underneath the stent antenna.

Far-field Radiation Characteristics
The 3D far-field radiation patterns were measured in the anechoic chamber at the Joint Laboratory for Microwave and Millimeter Wave Antenna Measurements at National Taiwan University of Science and Technology. Figure 13 shows the measurement setup. The stent antenna was mounted on a rotating stage and acted as a receiving antenna, and the probe antenna, which acted as a transmitting antenna, moved in the azimuth and elevation planes. The measured gain radiation patterns are shown in Figures 14 and 15. Note that the measured radiation patterns at 180 • were inaccurate due to the measurement setup where the transmitting antenna couldn't move to the position underneath the stent antenna.

Far-field Radiation Characteristics
The 3D far-field radiation patterns were measured in the anechoic chamber at the Joint Laboratory for Microwave and Millimeter Wave Antenna Measurements at National Taiwan University of Science and Technology. Figure 13 shows the measurement setup. The stent antenna was mounted on a rotating stage and acted as a receiving antenna, and the probe antenna, which acted as a transmitting antenna, moved in the azimuth and elevation planes. The measured gain radiation patterns are shown in Figures 14 and 15. Note that the measured radiation patterns at 180° were inaccurate due to the measurement setup where the transmitting antenna couldn't move to the position underneath the stent antenna.     Figure 14a,b shows the simulated and measured gain radiation patterns of the multi-connector stent antenna at a resonant frequency of 2.97 GHz in the E-and H-planes, respectively. Similarly, Figure 15a,b shows the simulated and measured gain radiation patterns of the single-connector stent antenna at a resonant frequency of 2.07 GHz in the E-and H-planes, respectively. It can be seen that the radiation efficiency of the single-connector stent antenna was superior to that of the multi-resonant stent antenna due to its approximate half-wavelength resonant mode shown in Table 2. Although the multi-connector stent had a larger gain compared to that of the single-connector stent, both stent-based antennas had donut-shaped radiation patterns with omnidirectional radiation in the azimuth plane, which would be suitable for arbitrary receiving angles. Figure 16 shows the gain and the radiation efficiency of the single-connector stent antenna at different frequencies. It was demonstrated that the highest radiation efficiency of the single-connector stent antenna was 74.5% at the resonant frequency of 2.07 GHz, with a corresponding gain of 1.38 dBi. It indicated that a single-connector stent could provide the maximum radiation efficiency and omnidirectional electromagnetic emissions for intravascular wireless communications.   Figure 14a,b shows the simulated and measured gain radiation patterns of the multi-connector stent antenna at a resonant frequency of 2.97 GHz in the E-and H-planes, respectively. Similarly, Figure 15a,b shows the simulated and measured gain radiation patterns of the single-connector stent antenna at a resonant frequency of 2.07 GHz in the E-and H-planes, respectively. It can be seen that the radiation efficiency of the single-connector stent antenna was superior to that of the multiresonant stent antenna due to its approximate half-wavelength resonant mode shown in Table 2. Although the multi-connector stent had a larger gain compared to that of the single-connector stent, both stent-based antennas had donut-shaped radiation patterns with omnidirectional radiation in the azimuth plane, which would be suitable for arbitrary receiving angles. Figure 16 shows the gain and the radiation efficiency of the single-connector stent antenna at different frequencies. It was demonstrated that the highest radiation efficiency of the single-connector stent antenna was 74.5% at the resonant frequency of 2.07 GHz, with a corresponding gain of 1.38 dBi. It indicated that a singleconnector stent could provide the maximum radiation efficiency and omnidirectional electromagnetic emissions for intravascular wireless communications.

Discussions
In this research, we proposed a single-connector stent antenna composed of the stent and SMA connectors for good radiation efficiencies. The SMA connector acted as a finite ground plane. The simulation results showed that the entire stent antenna had a half-wavelength current distribution shown in Figure 5 and a dipole-link radiation pattern shown in Figure 14. For all measurements conducted in this research, an RF choke was exploited to reduce the effects of the feed line and only considered the stent antenna. The measured radiation pattern demonstrated a dipole-link radiation pattern as expected shown in Figure 14. The discrepancy between the simulated and measured radiation patterns at 180° was due to the measuring equipment that the receiving antenna cannot move underneath the stent antenna and cannot measure the EM radiation at 180°.

Discussions
In this research, we proposed a single-connector stent antenna composed of the stent and SMA connectors for good radiation efficiencies. The SMA connector acted as a finite ground plane. The simulation results showed that the entire stent antenna had a half-wavelength current distribution shown in Figure 5 and a dipole-link radiation pattern shown in Figure 14. For all measurements conducted in this research, an RF choke was exploited to reduce the effects of the feed line and only considered the stent antenna. The measured radiation pattern demonstrated a dipole-link radiation pattern as expected shown in Figure 14. The discrepancy between the simulated and measured radiation patterns at 180 • was due to the measuring equipment that the receiving antenna cannot move underneath the stent antenna and cannot measure the EM radiation at 180 • .

Conclusions
In this paper, we present a stent antenna for intravascular monitoring and implantable wireless applications that employ the entire stent as an antenna. The main contribution of this work was the manipulation of the current distributions of the stent for good radiation performance with a short stent length by exploiting a single connector between each unit of the stent. Unlike a general stent, wherein each unit is connected with multiple connectors, our proposed stent can achieve good EM radiation, maintain good mechanical strength, and remain within fracture limits during balloon expansions. A prototype was designed, fabricated, and experimentally examined with an EM anechoic chamber. The fabricated stent antenna demonstrated an omnidirectional radiation pattern, a gain of 1.38 dBi, and a radiation efficiency of 74.5% at the resonant frequency of 2.07 GHz. The results of this research should contribute to the development of implantable wireless communications and intravascular monitoring of cardiovascular diseases.