A radial micro-vibration system for reducing friction of passive interventional guidewire

Cardiovascular and cerebrovascular diseases have become a serious threat to human health.¹ According to the report of the World Health Organization, cardiovascular disease (CVD) kills $17.9× 10 6$ people every year.² Vascular intervention surgery is widely used in the clinical treatment of CVD because of its advantages, such as a small incision and an accurate delivery of interventional instruments to the lesion,^3,4 reducing damage and improving treatment efficiency. In clinical vascular intervention surgery, doctors need to deliver a passive guidewire from the patient’s femoral artery (or radial artery) through a percutaneous puncture and then deliver it along the blood vessel to the lesion site,⁵ as shown in Fig. 1(a). Passive guidewires rely on manual operation, with the surgeon controlling the direction and position of the guidewire by inserting, rotating, and pulling it back. Therefore, the guidewire inevitably contacts the viscoelastic vascular lumen in multiple areas and is subject to resistance.^6,7 Especially, since most cardiovascular and cerebrovascular patients have complex malformations,⁸ such as narrow lumens and viscous blood, the guidewire is subject to time-varying strong resistance during delivery,^9–11 such as friction resistance,¹² viscosity resistance,¹³ and collision resistance.¹⁴ Time-varying strong resistance not only damages the blood vessel wall but also affects the delivery speed and accuracy, prolonging the operation time and increasing the surgical risk.^15,16

Surface modification of the guidewire is mainly a method used to reduce the resistance of interventional guidewires.^17–19 For example, Nurdin et al.²⁰ coated the catheter with a polyurethane layer to reduce the friction coefficient. Wang et al.¹⁸ proposed a soft guidewire coated with a 10–25  $μ$m thick hydrogel layer, whose surface friction was reduced by ten times and was biocompatible. Although surface modification can effectively reduce resistance and reduce damage to the blood vessel wall, the coating layer sometimes detaches from the surface of the guidewire. The reason is that due to the diameter of the guidewire being small and the coating thickness being thin, the coating layer is easy to suffer from varying strong stress under the disturbance of blood flow and repeated pushing, pulling, and twisting of the guidewire.^17,21 The coating fragments cause local tissue reactions, thrombosis, and other complications.^22–24 Moreover, the coating fragments are difficult to detect in surgery until complications arise. In addition, since smooth coating guidewires reduce controllability, risks, such as vascular compression and rebound puncture, are prone to occur.^25,26 Despite many studies that have improved the maneuverability of guidewires by designing active guidewires,^27–29 which combine advanced control technologies, force feedback, tactile feedback, and imaging navigation functions, most active guidewires are still in the laboratory stage and have not yet achieved mature clinical transformation. On the other hand, there are many types of guidewires used for intravascular treatment, and not all of them are able to be coated due to their functions. Therefore, vascular interventional surgery urgently needs a direct and effective method to reduce the resistance of clinical passive guidewires.

Based on the widespread application of ultrasonic vibration friction reduction characteristics in other fields,^30–32 our previous works applied ultrasonic microvibration to traditional passive guidewires to effectively reduce guidewire resistance.^13,33 Clinical passive guidewires do not require any special treatment, and their viscous resistance is reduced by 51.17% when lateral microvibration is applied.¹³ In addition, theoretical model analysis of microvibration verifies the effectiveness of applying radial microvibration to guidewires to reduce resistance.³³ However, previous works do not consider the complex vascular interventional surgical environment faced by the guidewire and the influence of multiple factors, such as vibration parameters, blood viscosity, and vascular malformations on the drag reduction effect. These factors generate the impact of a time-varying strong resistance. In addition, after working under high power for a long time of vibration, the clamping position of the piezoelectric ceramic sheet and the guidewire connection part are broken and fractured, which affects the experimental research and results. To provide a basis for accelerating clinical transformation and stable high energy microvibration, it is also necessary to optimize the design of the microvibration equipment.

In this paper, to extend our previous research, an ultrasonic microvibration-assisted guidewire drag reduction system is developed and used to study the interaction rule of multifactor on guidewire-assisted microvibration resistance reduction. We first design the sandwich structure of the ultrasonic transducer system and investigate the material selection and dimensions of each part, as shown in Fig. 1(b). Then, the step-type amplitude transformer is designed to drive the ultrasonic microvibration of guidewires and optimized by finite element simulation. The proposed system can apply radial ultrasonic microvibration with real-time resistance feedback to guidewires of various sizes. Subsequently, we further carry out one-factor and multi-factor analysis experiments on blood vessel models to investigate the individual effects on microvibration drag reduction. The ANOVA is used to analyze the significance of the factors and their interactions to obtain the guidewire microvibration-assisted drag reduction rules. Finally, the optimal combination of the factors for the drag reduction effect is demonstrated. The proposed guidewire drag reduction system and the multi-factor effect are of great significance for clinical interventional surgery applications.

In vascular interventional procedures, the friction generated by the guidewire movement in contact with the vessel wall is divided into the following three cases: static sliding friction, maximum static sliding friction, and sliding friction. According to the law of friction, the frictional resistance between the guidewire and the vessel wall can be reduced by decreasing the net friction time and contact area or the coefficient of friction. Based on the fact that ultrasonic vibration has friction reduction³⁴ and lubrication characteristics,^35,36 when one of the friction surfaces in the friction pair is in a state of ultrasonic vibration, the friction coefficient of the friction surface will be reduced.³⁰ Moreover, ultrasonic friction reduction is mainly due to the reduction of the effective contact area of the contact surfaces. The changing of the friction state between the contact surfaces changes the friction area of the contact surfaces.³⁷ 

This effect can be obtained by changing the friction state of the contact surface between the guidewire and the vessel wall by adding radial microvibration to the guidewire.³⁰ The guidewire applied with ultrasonic micro-amplitude vibration leads to a decrease in the coefficient of friction of the guidewire-vessel wall friction surface. It effectively reduces the contact area between the guidewire and the inner wall of the vessel, as shown in Fig. 1(c). As a result, the frictional resistance between the guidewire and the inner wall of the blood vessel is reduced, resulting in a decrease in resistance.³³ 

According to the microvibration analysis of the particle, applying ultrasonic vibration essentially reduces the frictional resistance of the guidewire by causing periodic intermittent contact between the guidewire’s whole length and the vascular wall. It indicates that the guidewire has a short contact time with the vascular wall and is separated for the rest of the time during each vibration cycle. As a result, the friction time and the contact area have both significantly decreased. Additionally, the presence of a vibration force causes the positive contact pressure to decrease. It is another reason that the guidewire applied with ultrasonic vibration could reduce the frictional resistance between the guidewire and the blood vessel wall. In addition, applying vibration changes the interaction form, which makes the interaction force between the surfaces intermittent. As a result, the dynamic friction coefficient between the two surfaces is reduced, and the static friction is converted into dynamic friction. Therefore, it reduces the frictional resistance of the guidewire.

Additionally, the levitation effect of the radiation pressure created by the ultrasonic vibration decreases the positive pressure, which reduces friction. Under ultrasonic vibration, the blood vessel and the longitudinal vibrating guidewire will generate a minimal suspension gap. The blood could then diffuse into this gap to form a lubricating film, which would further benefit the friction reduction. In summary, the coefficient of friction, the effective contact area, positive contact pressure, and the net friction time between the guidewire and the vessel wall are reduced after microvibration is applied to the guidewire. Therefore, the frictional resistance to the movement of the vibrating guidewire within the blood vessel will be reduced.

In our previous research,¹³ after multiple groups of experiments and long-term high-power operation, the clamping position of the piezoelectric ceramic slice and the fixed part of the guidewire had shattered and fractured in different degrees. This phenomenon mainly occurs at the position where the internal stress of the piezoelectric slice is at its maximum. It is that the mechanical limit of the piezoelectric ceramic slice is reached at the displacement node. Therefore, the single piezoelectric ceramic slice cannot be selected as a long-term vibration source for driving the guidewire ultrasonic microvibration.

This paper develops a new type drag reduction device for exerting ultrasonic micro-vibration on the guidewire. It consists of an ultrasonic transducer module, a guidewire resistance real-time detection module, and a guidewire motion control module. The ultrasonic transducer module includes an ultrasonic generator, a sandwich piezoelectric type transducer, and a stepped ultrasonic transformer.

We first designed the piezoelectric transducer of the ultrasonic system as a composite sandwich structure to drive the guidewire micro-vibration. Here, the axially polarized piezoelectric ceramic circle ring slice is used. The prestressed bolts and the matching insulating sleeves are used for connection. The application of a prestressing force can make the piezoelectric ceramic circle ring slice remain in a compression state all the time. Thus, it can resist the influence of strong vibration even under a high-power energy input. The structural diagram is shown in Fig. 3, which is mainly composed of three basic components: a metal front cover plate, a piezoelectric ceramic sheet stack, and a metal back cover plate. The piezoelectric transducer with a crystal stack structure is used. Therefore the polarization direction of the piezoelectric body, the direction of the external electric field, the combination mode, and the vibration mode of the crystal stack are all normal.⁴² The characteristics of signal conversion between piezoelectric ceramics are utilized to drive the guidewire to perform ultrasonic micro-amplitude vibration.

Subsequently, limited by the physical properties of piezoelectric ceramics, under the condition of high-power operation, the ceramic slice is easy to rupture due to the weak tensile strength and the inability to resist the energy input of high power; that is, the situation of the breakage of the piezoelectric ceramic slice occurred in the previous experiment of measuring the guidewire viscous resistance. Therefore, the sandwich transducer used in this paper applies prestress to the piezoelectric ceramic circle ring slice through metal blocks and prestress bolts to avoid the above-mentioned damage. This is mainly because the application of prestressing force can make the piezoelectric ceramic circle ring slice remain in a compression state all the time and resist the influence of strong vibration even under a high-power energy input, which is an improvement on the vibration source of the driving guidewire designed in our previous research. On the other hand, piezoelectric ceramics, as an insulating material, have poor thermal conductivity due to the restriction of their own material properties. Therefore, piezoelectric ceramics are extremely easy to generate heat under high power conditions. Moreover, the heat is difficult to transfer and is retained in the piezoelectric ceramics, which makes the internal temperature rise, resulting in a decrease in the overall energy conversion efficiency of the vibration source.

To apply high frequency and relative high amplitude vibration to the guidewire to study the method of friction reduction, we select the high-power piezoelectric ceramic PZT-8. The PZT-8 is high mechanical strength, easily polarized, and small dielectric loss. The excitation voltage on the thickness of the selected piezoelectric ceramic slice is $20 kV / cm$⁠, and the power capacity is $2.3 kW/( cm 3 kHz)$⁠. The piezoelectric effect is 20–30 times higher than that of a quartz crystal. With reference to the PZT-8 excitation voltage and the power capacity, the transducer piezoelectric oscillator selects eight pieces of ceramic disks with a thickness of 3.65 mm, a diameter of 15 mm, an inner bore diameter of 5.7 mm, and silver plating on both sides. Thus, the power of a piezoelectric transducer can calculate by $P=Powercapacity×V×f$⁠. $V$ is the volume of the transducer [ $V=πh( D o u t 2− D i n 2)/4$] and $f$ is the frequency. When they are superimposed to form a transducer, the positive pole is in the middle and the negative electrode is on both sides.

The front cover of the sandwich piezoelectric transducer plays the role of transmitting energy, ensuring that most of the energy in the transducer is efficiently radiated along its longitudinal front surface. Meanwhile, the front cover plate also has the function of impedance converter, which can effectively adjust the load impedance to keep the overall impedance value within the required range of piezoelectric ceramic elements. It improves the transmission efficiency of the transducer and ensures the frequency band width. The main function of the back cover plate of the transducer is to ensure that the energy generated by the transducer rarely radiates backward and transmit the energy to the transducer forward as much as possible, so as to reduce the energy loss.⁴² The addition of metal front and back cover plates to the sandwich piezoelectric transducer designed in this paper greatly improves its thermal conductivity, effectively solving the problem of poor heat dissipation in the mentioned vibration source. The thickness and transverse dimension of the metal material and the piezoelectric ceramic material greatly affect the temperature stability of the sandwich piezoelectric ceramic transducer. However, as long as a reasonable design is made, the temperature coefficient of the metal material elastic constant can compensate for the temperature coefficient of the piezoelectric ceramic material elastic constant so that the transducer can obtain a very small frequency temperature coefficient. Therefore, based on the designed calculation, the dimensions of the front and rear cover plates are obtained.

Based on the requirements, the aluminum alloy is used as the front cover material of the designed sandwich transducer, and the shape is laddering type. 45 steel is used as the back cover plate material, and the shape is cylindrical. In addition, it is required to improve the vibration velocity ratio of the front and rear cover plates. The impedance of the transducer and piezoelectric ceramics meets the following relationship: $Z F r o n t c o v e r< Z C e r a m i c< Z B a c k c o v e r$⁠. Then, the energy generated by the transducer is transmitted to the amplitude transformer in a single direction, and the maximum amplitude can be applied to the guidewire. According to the frequency [Eqs. (9) and (10)], the dimensions of the front and rear cover plates are obtained, as shown in Table I.

To meet the vibration amplitude of the experimental requirements, it is required to add the variable amplitude equipment to the vibration source system. The transducer needs to concentrate the unidirectional energy on a small area due to the geometric size limitation of the load guidewire. According to the calculation of input and output power load matching, the length and size of the amplitude transformer are determined. Due to the large ultrasonic amplitude and output resistance applied to the guidewire, the amplitude transformer driving the ultrasonic vibration of the guidewire is designed as a stepped type with a transition cross-sectional structure. This structure can satisfy the ultrasonic vibration under conditions of a large amplitude.

According to the relevant structural dimensions of the variator rod, we used ANSYS software to analyze the vibration mode of the step-type horn. Then, the specific performance parameters of the material of the ultrasonic horn, such as modulus of elasticity (⁠ $E=72GPT$⁠), density (⁠ $ρ=2800 kg / m 3$⁠), Poisson’s ratio (⁠ $μ=0.33$⁠), etc., were defined, respectively. The rear cover of the transducer is fixed on the designed support bracket, and the front cover is fully constrained with the piezoelectric ceramic solid connection. Therefore, the complete constraint is applied to the lower face of the transducer rod. The frequency interval is set from 5 to 50 kHz; the setup is set to extract the tenth order modes. As shown in Fig. 4, a total of six orders of modes are obtained within 5–50 kHz; only the third and sixth order vibration modes are longitudinal deformation, and the rest of the four cases are bending deformation. The third order vibration modes comply with the transducer requirements of its intrinsic frequency of 25 kHz operating frequency. 24 819 Hz is the intrinsic frequency of the sandwich-type transducer, as shown in Fig. 4(a). Figures 4(b)–4(f) show different intrinsic frequencies 6849, 24 819, 29 203, 29 651, and 37 676 Hz from the second to the sixth order, respectively. Furthermore, the resonance length $L$ and the nodal displacement $x 0$ of the ultrasonic horn are perfected, and the optimum value is found by calculating and analyzing the values several times. $L=74,75,76,77,78$ mm and $x 0=27,28,29,30,31$ mm were taken to carry out several modal analyses. The results indicate that when $L=77$ mm, $x 0=30$ mm, the analysis results and our theoretical design value are the closest.

To apply the power to the proposed sandwich-type piezoelectric transducer, the ultrasonic generator for driving the guidewire vibration is mainly composed of four parts: the waveform generating module, the power amplifier module, the impedance matching module, and the phase-locked loop module, as shown in Fig. 5(a). The main function of this device is to convert AC electrical signals into high-frequency electrical oscillation signals required by piezoelectric transducers. The XD2 general-purpose signal generator is selected as the ultrasonic signal source. The power source amplifies the AC220V/50 Hz $±$ 4 V sinusoidal voltage signal, converted into a frequency of 22.5 kHz $±$ 2%, which corresponds to the resonant frequency of the guidewire vibration driving equipment,³³ a peak-to-peak value of a 400 V high-voltage sinusoidal drive sandwich piezoelectric transducer. When the amplifier circuit works, the frequency and voltage of the output signal of the amplifier circuit are changed by adjusting the frequency and voltage of the signal generator within the range of (0–30) kHz and (0–4) V, and the oscilloscope detects the stable output sinusoidal voltage signal. Applying high-voltage sinusoidal AC voltage to the ultrasonic vibrator converts the ultrasonic electric oscillation into mechanical vibration, thereby meeting the design requirements for applying a variety of selectable frequencies and amplitudes to the guidewire.

Subsequently, the sandwich piezoelectric transducer was fixedly connected with the horn with pre-tightening screws. A through-hole is machined at the top of the amplitude transformer at the maximum amplitude, which can fix different types of guidewires. The end of the guidewire passes through this through-hole and is fastened using the upper bolt, which prevents the guidewire from slipping inside the through-hole, as shown in Fig. 5(b). This achieves applying radial micro-vibration to the proximal end of a conventional guidewire by using an ultrasonic generator to longitudinally vibrate a piezoelectric transducer. The equipment has a maximum input power of 50 W. To test the amplitude–frequency response curve, a piezoelectric ceramic plate was placed on top of the ultrasonic horn, close to where the guidewire was installed. Figure 5(c) shows the experimental result. Measurements have shown that the designed ultrasonic horn, fitted with a 1 Fr guidewire of 400 mm length, has a resonant frequency of 22.5 kHz $±$ 5%. It is capable of driving the guidewire to ultrasonic microvibration at various frequencies and amplitudes in the 0–40  $μ m$ range.

The guidewire vibration driving device was fixedly mounted on a bracket. Two linear sliding rails joined the bottom of the fixed bracket to the main bracket. On the left side of the main bracket, the microforce sensor (lrf400, FUTEK Advanced Sensor Technology, Inc., USA) with an accuracy of 0.03 mN was installed. The connector fixes the sensor to the fixed bracket, constituting the guidewire vibration driving force detection device. We measured the actual friction resistance of the guidewire in the artificial vascular model in real time using a 200 Hz sampling rate. Subsequently, we fixedly installed this part into the bottom linear slipway through the base. When the slide rail moved to the right, the friction force would be generated by the guidewire contacting the blood vessel. This friction force subsequently caused a micro displacement to the left in the fixed bracket on the slide rails, which transmit back to the sensor through pressure buildup in the connecting parts, as shown in Figs. 5(d) and 5(e). For the entire experiment, the friction force of the guidewire data was examined for real-time measurement.

The experimental medium was a physiological saline solution (⁠ $ρ=1.0 g / cm 3$⁠) and a simulated blood solution (⁠ $ρ=1.05 g / cm 3$⁠). The physiological saline solution and the glycerol solution (⁠ $ρ=1.263 g / cm 3$⁠) were mixed in the appropriate ratio to generate the simulated blood solution. With the comments of doctors and measurements of actual surgical data, the doctors must operate the guidewire in the blood vessels at low speed to ensure safety. The guidewire was, therefore, directly controlled by the linear slipway to move forward and backward at a constant speed of 1 mm/s.

In this case, we controlled the guidewire to stay on the same path to prevent collisions between other guidewire parts and the vascular wall. This was done to minimize the effect of other resistances on the sensor data and to confirm that the frictional resistance of the guidewire was generated by a contact between the guidewire and vascular wall. In the experiment, each parameter combination was repeated in ten experiments to reduce accidental errors. To improve the accuracy of the experimental results, we averaged the maximum values of all measured data to obtain the frictional resistance of the guidewire for analysis. To obtain the shape changes of the guidewire during the operation, a camera (⁠ $3264×2448$ pixels) is used to record the motion of the guidewire.

In addition, we used software (COMSOL 5.5) to simulate vibration of the micro-element of guidewire for analyzing the temperature change of blood caused by the vibration. In this case, the thermoviscous acoustic frequency domain module was suitable for Multiphysics interaction. The diameter of the guidewire particle and the blood vessel was set as 0.33 and 5 mm, respectively. The guidewire particle vibrates in the radial direction (Y axis). The thermoviscous boundary layer impedance condition is applied. The analytical solution is given by the Helmholtz decomposition of the acoustic particle velocity. Thus, the sound pressure can be obtained from $p(r,y)=iω ρ 0φ(r,y)$⁠, where $φ(r,y)$ is the velocity potential. The results show that the temperature is decreased for the center of the blood vessel to the edge. The simulation results indicate that the maximum temperature of the guidewire vibration is generated at the center of the vibration, as shown in Fig. 6(a). Moreover, the largest temperature of fluidity caused by guidewire vibrations of different frequencies and amplitudes is shown in Fig. 6(b). The temperature increases with frequency and amplitude increasing. According to the simulation results, the maximum temperature is not beyond 0.0026 K. It suggests that the temperature changes at a very small range when applying the vibration to a guidewire. Therefore, the heat generated by the vibration will not damage the blood vessels. Furthermore, according to the results in previous reports,⁴³ temperature measurements revealed an increase of less than $1 °$C in the vessel wall after 2 min of treatment at a power output of 5.5 W. Meanwhile, based on the pathological analysis, the applied vibration did not cause any damage to the endovascular lumen. Research conducted both in vitro and in vivo has demonstrated that therapeutic ultrasound can promote vasodilation and has no obvious harm to smooth muscle cells and endothelial cells. In other words, the damage to the blood vessel is unavoidable because of the mechanical collision in the blood vessels with a small size. The guidewire’s sliding speed affects the friction energy dissipation and coefficient of friction at the blood vessel–catheter interface.⁴⁴ 

Experiments were conducted to verify the effectiveness of the designed drag reduction system for reducing the frictional resistance of the guidewire in the blood vessel. The single-factor experiments were designed to study the influence of the vascular curve shape and the medium density on the drag reduction effect. In addition, multi-factor factorial experiments were designed to study the individual effects of vibration frequency, medium density, and simulated vascular bending shape on the drag reduction effect of the guidewire exerted with microvibration. Moreover, the interaction of multiple factors, the influence law, and the optimal drag reduction parameter combination were studied. Because the friction resistance mainly caused by the contact between the guidewire and the vascular wall and the effect of the simulated vascular diameter on the friction resistance is not obvious, the factor of a simulated vascular diameter is not considered.

To facilitate observation and obtain clear experimental images, the high-transparent medical soft pipes with an inner diameter of 5 mm were used instead of the simulated blood vessels, which is based on the clinical data of collaborators and the measurement of the EVE human body model in Fukuda Laboratory of Nagoya University, and their length was 500 mm. The maximum input voltage of the ultrasonic generator was controlled at 4 V. All the experiments measured the friction resistance between the guidewire and the inner wall of the simulated blood vessel in real time during the forward and backward movement of the guidewire in the simulated blood vessel. Under the conditions of each parameter combination, ten times of friction resistance measurements were also completed to reduce the accidental error.

The guidewire was applied with an ultrasonic vibration frequency of 22.5 kHz, which was compared with the frictional force on the guidewire without vibration. Experiments were carried out in a simulated blood vessel having a $180 °$ bend with the simulated blood solution. The blood vessel bending angles were also chosen based on the bending angle measured values of the blood vessel segment in the EVE human body model. The frictional resistances of the guidewire in the $180 °$ bending simulated blood vessel filled with the simulated blood solution under no vibration and at 22.5 kHz ultrasonic micro-vibration frequency are compared in Fig. 7. The red curves indicate the frictional resistance change of the guidewire in the simulated blood solution without vibration and represent the traditional surgery.

Figure 7(a) shows that the friction gradually increases initially with the guidewire advancement and then decreases again after reaching the maximum friction value at 18 s. According to the motion state of the guidewire in the blood vessel, it was obtained that as the slipway started to move forward, the guidewire was pushed to the inner wall of the blood vessel and started to contact the blood vessel wall. The guidewire was pushed to move forward and was subjected to sliding friction. After the slider moved forward for 18 s, the guidewire was thrust against the inner wall of the blood vessel and could not continue to move forward. At this time, the guidewire’s frictional resistance reached its highest value. The guidewire was subjected to the maximum static friction force. The guidewire was exerting the maximum force on the vascular wall at this point, maximally increasing the risk of causing harm to the blood vessel wall. Subsequently, the guidewire continued to be pushed forward, the static friction was converted to dynamic friction, and the guidewire continued to move along the vessel wall until it had fully advanced. During the backward process shown in Fig. 7(b), the guidewire was always subjected to sliding friction. The results indicate that the friction force was significantly lower than that of the non-vibrating guidewire at each moment when the vibration frequency of 22.5 kHz was applied, regardless of whether the process was forward or backward. Especially at 18 s, the frictional resistance of the guidewire with no vibration was 55.40 mN when moving forward. In contrast, the friction was 10.89 mN at a vibration frequency of 22.5 kHz (black line), indicating an 80.35% drag reduction. Similarly, during the backward process, the guidewire vibrating at 22.5 kHz exhibited a drag reduction rate of 82.48%.

In addition, to eliminate the lubrication effect of the fluid layer between the blood vessel wall and the guidewire, air was also selected as the experimental medium. Based on the experimental results in air, in contrast, applying ultrasonic vibration is the main reason for reducing the friction resistance of the guidewire. The above results are due to the applied micro-amplitude ultrasonic vibration converting the guidewire’s static friction into dynamic friction. Because the dynamic friction coefficient is lower than the static friction coefficient, the guidewire frictional resistance is accordingly reduced. Moreover, vibrating the guidewire greatly decreases the contact area and the net contact time between the guidewire and the vascular wall, significantly reducing the frictional resistance. The above experimental data showed that applying ultrasonic vibration at 22.5 kHz greatly reduced the frictional resistance of the guidewire compared to that without vibration. This corresponds to the guidewire vibration driving equipment’s resonance frequency. At this frequency, the guidewire vibrates the most obviously, and the vibration amplitude is the greatest. The experimental results show not only the vibration frequency but also the vibration amplitude could play a significant role in the frictional resistance of the guidewire. The drag reduction rate was also as high as 85.19% when the guidewire was advanced at 22.5 kHz and an amplitude of 40  $μ$m in a $180 °$ curved simulated blood vessel without solution, which is consistent with the experimental results in air during the backward process. The longitudinal vibrating guidewire and the blood vessel will form a minimal suspension gap between them under the ultrasonic vibration. The blood could then diffuse into this gap, forming a lubricating film that would be beneficial for friction reduction.

To further study the interactive influence of vibration frequency and amplitude applied by the designed drag reduction device on the friction resistance of the guidewire, different ultrasonic vibration frequencies were applied in the range of 20–25 kHz. The ultrasonic generator was controlled to input a maximum voltage of 4 V. The guidewire was driven by the different vibration frequencies in the $180 °$-bent simulated blood vessels filled with the simulated blood solution. Figure 8 shows the variations in the average maximum frictional resistance $F$ during the forward and backward movement processes of the guidewire and the statistical diagram of the drag reduction rates, which are calculated by comparing $F$ with the frictional resistance without vibration $F 0$⁠.

During the advancement process, the maximum frictional forces of the guidewire with ultrasonic vibration were less than that without micro-vibration at all tested frequencies [Fig. 8(a)]. The frictional resistance was the lowest at 22.5 kHz, and the trend of the friction reduction curve was similar to the amplitude–frequency response curve of the designed guidewire vibration driving equipment. In other words, the drag reduction effect varied with the amplitude–frequency response curve. The same trend was observed when the guidewire moved backward [Fig. 8(b)]. This is mainly because the driving equipment reached resonance at about 22.5 kHz, which maximized the amplitude. The amplitude transformer drove the guidewire to obtain the maximum vibration energy and displacement. The vibration increases with the vibration energy. At high frequencies, the guidewire was far from the inner wall of the simulated blood vessel. The friction coefficient and the friction time between the guidewire and the inner vascular wall accordingly decreased, which, in turn, reduced the average friction.

It also indicates that both the vibration frequency and the vibration amplitude play a significant role in the frictional resistance of the guidewire. The different vibration devices with different sizes of guidewire exhibit different resonance frequencies. As the vibration device is further updated, the resonance frequency of the device will change accordingly. Hence, it should figure out the corresponding optimal resonance frequency to obtain the best antifriction effect. However, when other vibration frequencies far from the resonant frequency, such as 20 and 25 kHz, are applied, the experimental device may experience mechanical vibrations, which can affect the reading of the mechanical sensor and cause experimental data fluctuation. The drag reduction effect is not obvious.

The above experimental results demonstrated that the designed friction reduction device can effectively reduce the friction between the guidewire and the simulated vascular wall, and the friction reduction is most obvious at the resonance frequency of the ultrasonic transducer. The drag reduction rate varied with the change in the amplitude–frequency response curve.

Blood vessels are always curved in reality. For example, in coronary arteries and aortic arch vessels, bifurcation vessel angles and vessel bending angles of about $30 °$⁠, $45 °$⁠, $60 °$⁠, $90 °$⁠, and $180 °$ are common, which are confirmed from the EVE human body model. According to the report from the doctor’s clinical operations, the friction resistances will also vary greatly when the guidewire moves in the blood vessels with different shapes and bending angles. Therefore, the blood vessel bending shape is also an important factor affecting the friction resistance between the guidewire and the blood vessel wall. Based on different vascular distribution patterns,⁴⁵ six simulated vascular shapes with different bending angles were designed for comparative experiments, as shown in Fig. 9. Then, the frictional resistance values of the guidewire in simulated blood vessels with different bending angles were measured with or without ultrasonic micro-amplitude vibration applied to examine the influence of the vascular bending shape on the frictional resistance.

In the single-factor experiment of the influence of the blood vessel’s curved shape on the frictional resistance of the guidewire, the vibration amplitude was 40  $μ$m, and the vibration frequencies were 0 Hz and 22.5 kHz. The simulated blood vessel shapes were set to $30 °$⁠, $45 °$⁠, $90 °$⁠, $180 °$⁠, and $360 °$ bends, totaling 5 simulated blood vessel curved angles. The comparison of the average maximum friction value of the guidewire under different simulated vessel bending angles is shown in Fig. 10. Figure 10(a) shows that in the forward process, when the guidewire is not applied with vibration (the red and blue curves), the friction resistance value of the guidewire increases with the increase of the bending angle. However, when the vibration frequency of 22.5 kHz is applied, the friction resistance of the guidewire is significantly reduced in both the simulated blood solution and the normal saline solution (the black and green curves). In simulated blood, the friction reduction rate of the guidewire applied with the resonance frequency in simulated blood vessels with 5 bending angles of $30 °$⁠, $45 °$⁠, $90 °$⁠, $180 °$⁠, and $360 °$ is 84.88%, 85.19%, 56.7%, 80.35%, and 77.40%, respectively. In a physiological saline solution, the friction reduction rates are 78.6%, 83.77%, 56.66%, 78.55%, and 74.32%, respectively. The average friction reduction rates have reached about 80%, which proves that the designed setup has an obvious friction reduction effect on curved blood vessels. The reason is that the guidewire moves close to the blood vessel wall when it is in the bending position, and the friction resistance of the guidewire body will become greater. After radial micro-vibration is applied to the guidewire, the guidewire moves away from the inner wall of the blood vessel at a high frequency. The friction coefficient and the friction time between the guidewire and the inner wall of the blood vessel are correspondingly reduced. The average friction value is reduced accordingly, and the drag reduction effect is obvious.

The analysis of the above single-factor experimental results also showed that the change in the vascular bending angle has an impact on the frictional force of the guidewire. As the density of the medium increases, the frictional resistance increases slightly. However, the change in the medium density has no effect on the drag reduction rate. The same trend was observed when the guidewire moved backward [Fig. 10(b)]. However, it is observed that the drag reduction effect of the guidewire in the $90 °$⁠, $180 °$⁠, and $360 °$ bends is not as good as that in the $30 °$ and $45 °$ bends. Therefore, these three simulated vessel shapes with simulated vessel angles of $90 °$⁠, $180 °$⁠, and $360 °$ were further studied in the simulated blood solution. A total of seven vibration frequencies in the vibration frequency range of 0 Hz to 25 kHz were applied to the moving guidewire to verify the effectiveness of the designed drag reduction equipment in reducing the friction resistance of the guidewire in the blood vessel with a large bending angle.

Subsequently, we further explore the interaction between the bending angle and the vibration frequency and amplitude. Figures 11(a), 11(c), and 11(e) show the movement position and state of the guidewire in a simulated blood vessel at $90 °$⁠, $180 °$ and $360 °$ bent, respectively. In the simulated blood solution, the change in the average maximum friction value and the friction reduction rate in the process of the guidewire moving forward and backward driven by different microvibration frequencies are shown in Figs. 11(b), 11(d), and 11(f). The results show that in the blood vessels of these three bending angles, the frictional force on the guidewire is the smallest with the vibration frequency applied from 22 to 23 kHz. The drag reduction rate also varies with the change of the amplitude–frequency response curve. However, in the $90 °$ vessel bend, the friction reduction rate during the guidewire advancement is significantly lower than that of the guidewire in other curved angles of blood vessels, which was 59.9% [Fig. 11(b1)]. The drag reduction trend when the guidewire is moved backward [Fig. 11(b2)] is the same as that when the guidewire is moved forward, but the maximum drag reduction rate further decreases. The reason should be that, under the influence of the simulated blood vessel shape of a $90 °$ bend, the vibration energy of the guidewire will be attenuated in the bending area because the guidewire enters the simulated blood vessel in the bending state, resulting in the vibration state of the guidewire being weakened compared with other bending angles. However, the guidewire applied micro-amplitude ultrasonic vibration still has a good drag reduction effect in this kind of vascular shape. The experimental data show that the average maximum friction value of the guidewire in the forward process is almost twice that in the backward process under the condition of no vibration. This also shows that in the process of pushing a guidewire into a similar blood vessel, the operation is riskier, and the operation method of the guidewire with a safety guarantee is needed. When the guidewire is applied with 22.5 kHz micro-amplitude ultrasonic vibration, the friction force of the guidewire can be controlled at about 19 mN in both forward and backward states, reducing the hazardous risk resistance by half. It is of great significance to reduce the risk caused by the guidewire passing through such blood vessels.

In the simulated blood vessel shape with a $180 °$ bend, the maximum friction force of the guidewire without vibration reached 55.40 mN, which was 6 mN larger than that when the simulated blood vessel had a $90 °$ bend. It can be seen that with the increase in the bend of blood vessels, the operation risk increases gradually. However, after advancing the guidewire with the addition of microvibration, the frictional force on the guidewire was significantly reduced to 14.06 mN, thus demonstrating that the addition of ultrasound microvibration is still effective in reducing the hazardous frictional resistance of guidewires during this type of intravascular movement. The movement of the guidewire in the simulated blood vessel at a $360 °$ bending angle is depicted in Fig. 11(e), as is the contact with the inner wall of the simulated blood vessel during the movement. When starting to push the guidewire, the guidewire does not move forward at first but first expands to the inner wall of the blood vessel. The contact area with the inner wall of the blood vessel gradually increases and enters the midway state. Then, the guidewire begins to move along the inner wall of the blood vessel to the termination position. The results are consistent with those of the simulated blood vessels at $90 °$ and $180 °$ bending angles, and the frictional force of the guidewire has the minimum value in the range of 22–23 kHz. When no vibration is applied, the maximum friction resistance of the guidewire moving forward in the $360 °$ bend is 62.30 mN. Compared with the force in the $90 °$ and $180 °$ bends, the risk of injury to the simulated inner wall of the blood vessel is also the largest. However, when 22.5 kHz vibration is applied to the guidewire, the friction force is reduced to 13.24 mN, and the friction reduction rate can still reach 77.41%.

The above results show that whether the guidewire moves forward or backward, the exerted microvibration has a good effect on reducing the friction resistance of the guidewire in the complex shape of blood vessels. Moreover, the larger the vascular bending angle, the greater the friction resistance of the guidewire, but the drag reduction effect is still good. This greatly improves the smoothness and safety of the guidewire operation. However, there are also some bending angles that may cause attenuation of the vibration energy of the guidewire, resulting in a negative impact on the drag reduction effect. Additionally, the change of the medium density has little effect on the drag reduction effect of the guidewire supplemented with ultrasonic micro-vibration.

According to the above studies of the guidewire frictional resistance, it is found that the presence or absence of applied vibration, the applied vibration frequency, the vibration amplitude, the medium density, and the vascular bending angle all have an impact on the frictional resistance of the guidewire. There is a law that the bending degree of the blood vessel affects the friction reducing effect of the vibration frequency and the vibration amplitude. According to previous studies,³³ the vibration amplitude is the most significant factor and has an interactive relationship with the simulated blood vessel bending angle but has little correlation with the experimental medium. Therefore, the analytical factorial experimental design was chosen with the aim of comprehensively analyzing the main effects and interactions of the vibration frequency, the simulated blood vessel curved shape, and media density on guidewire friction resistance by means of an analysis of variance (ANOVA) method and SPSS software. The friction reduction law of the microvibration-assisted method on the guidewire is explored. The optimal combination of factors that can obtain the best resistance reduction effect is obtained.

An artificial blood vessel model with an inner diameter of 5 mm and a length of 450 mm was selected. The ultrasonic generator was controlled to input a maximum voltage of 4 V. Each factor is set at multiple levels. The vibration frequency is set as 0 Hz, 20 kHz, 20.5 kHz, 21 kHz, 21.5 kHz, 22 kHz, 22.5 kHz, 23 kHz, 23.5 kHz, 24 kHz, 24.5 kHz, and 25 kHz, totaling 12 levels; the simulated vessel shape is set as $30 °$⁠, $45 °$⁠, $90 °$⁠, $180 °$⁠, and $360 °$ bends, totaling 5 levels; the medium density is set as 1.0 and 1.05 g/cm $3$⁠, totaling 2 levels. Comprehensive experiments of the above-mentioned multi-factors were carried out, and all combinations of each factor and level were experimentally explored. The results of the variance analysis are listed in Table II. Among the influencing factors, the effect of the vibration frequency and the simulated vascular bending angle on the drag reduction effect are significant factors. According to the $F$ value, the factor simulated blood vessel bending angle is the most significant one, followed by the factor vibration frequency. The medium density is a non-significant factor. Meanwhile, the interaction of the three influencing factors was analyzed as Tables III–V.

The results show that there is only an interaction between the vibration frequency and the simulated blood vessel shape. There is no interaction between the medium density and the simulated blood vessel shape, or between the vibration frequency and the medium density. In this case, since the friction resistance value of the guidewire is as small as possible, it is used as an experimental index. The optimal drag reduction scheme should select the level corresponding to the minimum value of $K$ among the factors of the vibration frequency, the vibration amplitude, the medium density, and the simulated blood vessel shape. That is, based on the results of the multi-factor factorial experiments, the factor vibration amplitude takes the maximum value, the vibration frequency takes the seventh level, and the simulation blood vessel bending angle takes the first factor. The interaction between the vibration frequency and the blood vessel bending angle and the interaction between the vibration amplitude and the blood vessel bending angle are also significant. The measurement value is the smallest. Finally, the optimal combination of drag reduction parameters was obtained in a simulated blood vessel with a bending angle of $30 °$⁠, the exerted vibration frequency was 22.5 kHz, and the vibration amplitude was 40  $μ$m. The measurement result of the friction resistance of the guidewire in the simulated blood solution was 3.34 mN, which is the smallest value among all experimental results. The friction compared with the frictional resistance value of 21.25 mN without vibration under the same conditions. It still has a good anti-friction effect in the simulated blood vessels at other bending angles.

Overall, during the process of applying ultrasonic vibration to the guidewire, the frictional resistance between the guidewire and the inner wall of the blood vessel is determined by the comprehensive action of the applied vibration amplitude, the vibration frequency, and the simulated blood vessel bending angle. The ideal friction resistance value of the guidewire can be obtained by selecting the appropriate parameter combination. As shown in Fig. 12, the response surface curves of the frictional resistance in the process of the forward and backward movements of the guidewire with or without ultrasonic vibration are constructed by the vibration frequency and the simulated vascular bending angle in the simulated blood solution and the physiological saline solution, respectively. Here, the vibration amplitude is 40  $μ$m. According to the response surfaces, the frictional resistance of the guidewire after applying ultrasonic microvibration shows a significant downward trend compared with the traditional guidewire operation. When the guidewire moves forward in physiological saline and simulated blood solutions, the “optimal drag reduction zone” is located at the simulated blood vessel bending angle of $45 °$⁠. It occurs at the resonance frequency of the ultrasonic generator driving the guidewire vibration, where the vibration amplitude of the guidewire is the largest. When the guidewire moves backward, the “optimal resistance reduction zone” is located at the simulated vessel bending angles of $45 °$ and $360 °$⁠. It also appears at the resonance frequency of the ultrasonic generator driving the guidewire vibration and the maximum guidewire vibration amplitude.

The above results indicate that when using the applied ultrasonic vibration setup to reduce the guidewire friction resistance, the resonance frequency of the ultrasonic generator and the maximum vibration amplitude are used to drive the guidewire to vibrate. It can obtain the maximum friction reduction rate, that is, the minimum friction resistance value of the guidewire. In addition, according to actual needs, other vibration frequencies and amplitudes can be selected to drive different types of guidewire to reduce the frictional resistance of the guidewire to a suitable range.

In this paper, a novel drag reduction device for guidewire ultrasonic micro-amplitude vibration is developed to study multifactorial interaction. First, the vibrating source for driving the guidewire’s micro-vibration is optimized. To avoid damage to the piezoelectric ceramics by high-power ultrasonic vibration, the piezoelectric transducer is designed as a composite sandwich structure with a preload. We set the structural speed ratio and the amplitude lever to apply the transducer energy to the guidewire with a maximum amplitude in a single direction. Then, a real-time force feedback drag reduction system for guidewire operation was designed. Subsequently, we analyzed the main effects and interactions of the factors affecting drag reduction during micro-vibration operation by simulating the vascular and blood environments. The system effectively reduced the frictional force of the guidewire by applying microvibration. The results showed that the resistance reduction rate varied with the amplitude–frequency response curve. The damping effect is still obvious even when the blood vessel curves significantly. The multifactorial experiments also showed that the vibration frequency, the amplitude, and the simulated vessel bending angle were the most important factors,⁴⁶ followed by the vibration frequency, while the medium density was a non-significant factor. Only the vibration frequency and amplitude both interacted with the simulated vessel shape. Finally, the results show that driving the guidewire at the maximum amplitude at the resonant frequency of the ultrasonic transducer reaches the maximum friction reduction (optimal drag reduction zone). Depending on the actual situation, the designed system can utilize different parameter combinations to reduce the guidewire’s frictional resistance to an appropriate range.

Our system will significantly explore the effect of the sliding speed of the guidewire in the future. Future work will focus on experimental studies in real blood vessels to explore the influence of other factors on the proposed method, such as the guidewire size and speed and thermal effects. In summary, this study demonstrates that the developed guidewire microvibration-assisted drag reduction system effectively reduces intravascular guidewire frictional resistance, revealing important drag reduction factor rules. It has potential applications in assisting surgeons to perform surgeries more smoothly.

This work was supported by the Youth Research Program of the Open University of China (No. Q23A0021), the National Natural Science Foundation of China (No. 62003128), and the National Natural Science Foundation of China (No. 62103299).

The authors have no conflicts to disclose.

Chaonan Zhang: Conceptualization (lead); Data curation (lead); Methodology (lead); Validation (lead); Visualization (lead); Writing – original draft (lead); Writing – review & editing (lead). Liping Pan: Validation (equal); Writing – review & editing (equal). Xiajing Wang: Methodology (equal); Validation (equal). Ru Guo: Investigation (equal); Methodology (equal). Yan Zhao: Supervision (equal); Writing – review & editing (equal). Shuzhang Liang: Conceptualization (equal); Investigation (equal); Methodology (equal); Supervision (equal); Writing – review & editing (equal).

The data that support the findings of this study are available from the corresponding author upon reasonable request.