Magnetically controlled microrobotic system for programmable stiffness tuning and active steering of microcatheters

Domain: physics
Published: 2025-12-23
Editorial outcome: minor_revision

Abstract
Surgical tasks in small tortuous lumens demand interventional instruments with controllable mechanical adaptability. However, current microcatheters lack a non-disruptive, integration-ready strategy for dynamic stiffness tuning—critical for meeting the divergent mechanical demands for compliant steering and stable advancement. Here, we present a microrobotic system based on a helix-shaped magnetic soft microrobot (Helixoft) that compatibly integrates with commercial microcatheters (down to 300 μm in diameter), enabling continuous stiffness tuning (up to 40-fold) and precise steering, both controlled magnetically, free of any other potentially harmful stimuli. Stiffness tuning and active steering are independently controlled via a decoupled control strategy by the helical motion and torque-driven bending of independent microrobot components. This stiffness and structure reconfiguration allow the integrated microcatheter to perform large-angle navigation, precision payload delivery, and localised tissue biopsy without unintended buckling or tissue damage. We validate the system in both ex-vivo oviduct biopsy and in-vivo drug delivery to the fourth-generation bronchi of live pigs. The Helixoft system provides a minimally disruptive robotic strategy for the mechanical reconfiguration in confined and sensitive luminal environments.

Introduction
Minimally invasive surgery (MIS) has significantly advanced modern clinical practice, offering reduced trauma, fast recovery, and minimal scarring, particularly for interventions within small lumens, such as blood vessels, bronchi, and oviducts1,2. Microcatheters are indispensable tools for these procedures3–5, enabling precise diagnosis and treatment interventions, such as biopsy6–8, stent placement9,10 and reagent delivery11–13. For performing complex tasks in intricate and tortuous lumens, microcatheters must achieve precise control over their steerability and compliance adaptability, for effectively manoeuvring through complex paths while maintaining sufficient mechanical stability for performing tissue interactions without causing unintended damage.

Previous efforts to enhance steerability have led to the development of surgical catheters based on tendon-driven14–17 and concentric tube structures18,19, as well as conductive20–22 and magnetic shape-transformable materials11,23–25. Among these devices, magnetically controlled soft catheters have demonstrated distinct advantages, offering precise steerability through the highly penetrable and biosafe actuation mechanism afforded by the magnetic field11,23,26. Various magnetic catheters, featuring functionalities such as active navigation27, light delivery28, thrombus removal29,30, and force sensing31,32, were developed by Nelson et al., Zhao et al., our group, and others. Such magnetic catheters, aligned by the magnetic torque, allow for multi-degree-of-freedom orientation control within absolute spatial coordinates, without the need for antagonistic tendons.

However, most magnetically controlled catheters have been designed with constant stiffness, limiting their adaptability to the dynamic mechanical demands of complex operations. In clinical practice, an effective catheter is desired to remain soft during steering yet sufficiently stiff during advancement33. The soft state facilitates smooth navigation through tortuous lumens, whereas the stiff state prevents buckling during advancement in straight luminal segments. Additionally, catheters equipped with end-effectors, such as biopsy needles34 and graspers11, also need such switchable stiffness states to realise both safe navigation around sensitive tissues and efficient penetration at target sites during procedural tasks35,36. Accordingly, both the navigational and operational performances of microcatheters fundamentally depend on achieving reliable and controllable stiffness tuning.

Stiffness tuning strategies previously reported focused primarily on two approaches: material-based and structure-based strategies. Material-based strategies employed viscosity-tuneable materials, such as magnetorheological37–39 and electrorheological fluids40–42, or phase-changing materials including shape-memory polymers20,43,44 and low-melting-point alloys12,33,45,46. These materials enabled stiffness tuning in response to magnetic, electrical, or thermal fields. However, their reliance on such intense stimuli raised considerable safety concerns for delicate biological environments, including risks of toxicity, thermal injury, and electrical damage. Additionally, integrating stimulus-delivery elements, such as heaters or electrodes, faces significant fabrication challenges, particularly when miniaturising the microcatheters for applications in small lumens.

Structure-based strategies, such as joint stacking47,48, particle jamming49, layer jamming50, fibre jamming51, sliding sleeve, and concentric tube19,52, were proposed by Yang et al., Zhang et al., and others as more biocompatible alternatives for modulating catheter stiffness. Although these structure-based strategies mitigated safety concerns, such as thermal risks, they involved complex integration processes and inevitably led to increased device size, restricting their applicability in small and delicate anatomical environments.

Moreover, previous catheters based on material- and structure-based strategies were primarily developed with diameters exceeding 2 mm, a compromise size requirement to accommodate bulky actuation components, which in turn restrict their applicability mainly to cardiac and gastrointestinal interventions. To date, downsizing stiffness-variable catheters to diameters approaching 1 mm—a critical threshold for vascular and reproductive system applications—remains a significant challenge3,28,53.

To address these challenges, we present a microrobotic system comprising a helix-shaped magnetic soft microrobot (Helixoft) that seamlessly integrates with commercial microcatheters, to enable continuous stiffness tuning and arbitrary-directional steering, thereby achieving adaptive navigation and controllable tissue penetrability (Fig. 1). The Helixoft microrobot integrates both magnetic responsiveness and mechanical compliance through its constituent components, a rigid magnetic helix and a soft microtube, enabling decoupled control over the stiffness tuning and steering of the integrated microcatheter, based on the material magnetisation–device structure–field pattern triad design. Integrated onto microcatheters with scalable diameters down to 300 μm, the Helixoft enables continuous stiffness tuning across a 40-fold range, superior to previous structure-based strategies that achieve only binary or discrete modulation states. The stiffness tuning and steering processes are solely controlled by low-intensity magnetic fields with no need for thermal, optical, or electrical stimuli that compromise tissue safety or microcatheter integrity. The Helixoft-integrated microcatheter can be also attached with auxiliary microsurgical tools, such as mini-cameras, electrodes, and laser fibres, thereby enabling multimodal surgical functionalities, such as on-board imaging or ablation. Supplementary Table 2 summarises the performance of the Helixoft-integrated microcatheter in comparison with representative variable-stiffness and steerable catheters.

To quantitatively predict the pose of the Helixoft-integrated microcatheter, we have developed a mechanical model capable of accurately describing the microcatheter’s mechanical behaviour. We establish both the precise numerical result and the approximate analytical expression for the trajectory of the microcatheter’s pose, providing a theoretical basis for precise control during operation. Combined with the robotic catheter manipulator, magnetic control and imaging modules, the Helixoft-integrated microcatheter performs active navigation, precision payload delivery, and localised tissue biopsy within complex luminal networks, as validated through in-vivo bronchial interventions and ex-vivo oviduct biopsy. Together, the Helixoft microrobotic system offers a generalisable control strategy with minimal tissue damage for the steerability and stiffness tuning of microsurgical instruments operation within small and sensitive anatomical environments (e.g., bronchioles, cerebral vasculature, and fallopian tube).

Results
Design and magnetic actuation principle of the Helixoft microrobot
The Helixoft microrobot comprises a rigid magnetic helix, a soft microtube, and a tubular distal end (Fig. 1). The rigid magnetic helix and tubular end are composed of medical-grade stainless steel, whereas the soft microtube is moulded from biocompatible silicone elastomer (Supplementary Fig. 1). This combination of materials ensures both compliance and biocompatibility for surgical applications. A cylindrical anchor is affixed to the soft microtube’s outer layer between the rigid magnetic helix threads to prevent unintended translational sliding of the rigid helix caused by frictional forces from the surrounding tissue. To further reduce friction during navigation, the entire microrobot is coated with a biocompatible lubricating layer. The tubular end, currently fabricated from a commercial biopsy needle for demonstration purposes, can be replaced with alternative biocompatible magnetic materials depending on specific purposes. The open architecture and dimensional scalability of Helixoft allow for seamless integration onto commercial microcatheters with diameters as small as 300 μm via simple adhesive bonding.

Continuous stiffness tuning of the Helixoft-integrated microcatheter is achieved by dynamically rotating the helix to adjust the length of the soft section of the Helixoft under a rotating magnetic field (R-field) through programming the rotation angle (Fig. 1). As the R-field rotates, the high-stiffness magnetic helix undergoes a screw-like helical motion, gradually overlapping or uncovering the soft microtube, thereby tuning the overall stiffness of the microcatheter. Steering of the Helixoft-integrated microcatheter is achieved by aligning its magnetic end under a static magnetic field (S-field) through programming the field direction. For clinical safety, the system is designed to operate in a sequential manner across distinct modes, including stiffness tuning, advancement, steering, and biopsy. Before each operation, the microcatheter stiffness is tuned under the R-field to the appropriate state. During advancement through straight luminal segments, the catheter is set to a high-stiffness state to prevent distal buckling. When approaching bifurcations or traversing tortuous lumens, its stiffness is reduced under the R-field so that the magnetic distal end can bend the soft microtube under the S-field for active steering. Upon reaching the target site, the microcatheter is switched back to a high-stiffness state to provide sufficient mechanical stability for biopsy or other manipulations. In practice, the Helixoft-integrated microcatheter can be set to three predefined states, simplifying its use for standard minimally invasive surgical scenarios without the need for complex continuous tuning. The low-stiffness state enables high flexibility for large-angle steering under the S-field, while the high-stiffness state ensures structural rigidity for stable advancement and tissue penetration. The medium-stiffness state provides a mechanical response, adaptable to varied environmental conditions.

Mechanical modelling and pose prediction of the Helixoft-integrated microcatheter
Modelling the stiffness tuning and magnetic steering is critical for enabling efficient control of the navigational and surgical performance of the Helixoft-integrated microcatheter (Fig. 2a). After integration, the microcatheter’s average stiffness is dependent on the overlapping length of the rigid helix and the soft microtube, and thus tuneable through the rotation angle of the R-field. Afterwards, with a given stiffness, the microcatheter bending (orientation) is dependent on the magnetic torque, and thus adjustable through the field direction of the S-field. Based on an average constant curvature model, the microcatheter average stiffness satisfies EI=L1+L2(L1/E1+L2/E2)I (see details in Supplementary Notes 1 and 2)54. The pose of the microcatheter’s end is established as:

sin(θr/θ) = LM EI B / p
x = Lsinθ + (LG + LM)cosθ
y = L(1 - cosθ) + (LG + LM)sinθ

We achieve the microcatheter pose through both an analytical solution after a certain approximation and a numerical solution by solving the governing system of nonlinear equations. During modelling and measurement, the results are dependent on the microcatheter size. For demonstration purposes, we focus on commercial microcatheters with diameters ranging from 300 μm to 1 mm, providing a representative range of sizes and stiffness levels (Fig. 2b). The numerical analysis is conducted using a hybrid optimisation method combining a geneticalgorithm for global search and a local optimisation strategy. The approximate analytical expression is given as:

θ = arccot(1/sinr * (cosr + E1E2 / (L1E2 + L2E1) * I / MB)), when θ < π
θ = MB / (2r * (L1E1 + L2E2) / (L1E1 + L2E2) + MB) - 1, when θ > r - π

Methodological details and results are shown in Supplementary Note 2 and Fig. 2c.

Decoupled control for stiffness tuning and active steering of the microcatheter
The Helixoft system enables decoupled control between the stiffness tuning and active steering. Supplementary Movies 1 and 2 demonstrate their decoupled motion, enabling sequential and independent control of both functions through magnetic field programming. Specifically, the radially magnetised rigid helix responds to the R-field (3 mT), producing screw-like helical motion with a translational speed of up to 5 mm/s to tune the microcatheter’s stiffness (Fig. 2d). Steering is subsequently achieved by aligning the radially magnetised end under the S-field. The bending response depends on the preprogrammed stiffness level. As shown in Fig. 2f, a 1 mm-diameter microcatheter achieves bending angles ranging from 24° at the high-stiffness state (L/L=0) to 73° at the low-stiffness state (L/L=1), when exposed to an S-field perpendicular to the distal end magnetisation. Adjusting the S-field direction further enables precise control of the bending angle, reaching up to 118° at the low-stiffness state under a 150-mT S-field of oriented at 150° (Fig. 2e).

To avoid unintended bending of the microcatheter during helical motion, the magnetic distal end is also radially magnetised. In contrast, an axially magnetised distal end would bend and rotate conically under the R-field, potentially causing tissue damage through friction and compression. The present decoupling control strategy, based on the material magnetisation–device structure–field pattern triad, enables fully programmable and simultaneous regulation of both motion and mechanical properties, promising a generalisable decoupling control approach for magnetic soft robotics systems.

Control accuracy of the stiffness tuning and the microcatheter pose
The prediction accuracy of the developed model on the average stiffness and the distal end’s pose is validated by a standard three-point bending test and real-time tracking, respectively. As shown in Fig. 2c and Supplementary Figs. 2–5, both the approximate and numerical solutions exhibit strong agreement with the experimental results. Stiffness tuning results deviate across microcatheter sizes. For instance, among the three demonstrated microcatheters, the 600-μm microcatheter achieves the highest stiffness tuning range over 40-fold (36 kPa–1500 kPa), whereas the 1-mm microcatheter achieves the finest tuning precision during rotation (0.12 kPa/deg). For a given microcatheter, the magnetic end orientation increases almost linearly with the S-field strength, while increasing the R-field rotation angle progressively stiffens the microcatheter, reducing its bending response. As stiffness rises, the microcatheter’s bending transitions from a pure-rod behaviour toward a cantilever-like elasticity, deviating from linearity.

Comparing the measured values of the microcatheter poses with the numerical and simplified analytical solutions, we assess our model’s prediction accuracy by a classic mean absolute percentage error (MAPE) value test, in which the MAPE value represents the average deviation of the model’s predictions from the actual values (Supplementary Fig. 6). Generally, a MAPE below 10% suggests highly accurate forecasting, while 10–20% indicates good accuracy55,56. For the distal end’s pose coordinates of the microcatheter (x, y, θ) when the R-field rotates 0, 3 and 6 turns, the MAPE values between the numerical model and the experimental measurement are 4.7%, 5.1%, and 8.5%, respectively. Based on our results, the model provides a highly accurate prediction of the microcatheter pose (≥91.5%), especially at the low stiffness range (33 kPa–290 kPa).

Distributed stiffness control with a dual-segment module
To demonstrate the scalability of our stiffness tuning mechanism, we have implemented a multi-segment design that enables multi-level stiffness tuning across different regions of a single microcatheter57. As shown in Fig. 3, a dual-segment stiffness tuning module is integrated along the microcatheter. Each segment consists of a rigid magnetic helix coupled to a soft microtube. Independent control of each helix is achieved via the interplay between the magnetic actuation and local friction. A permanent cylindrical magnet (1.5 cm in diameter) generates a rotating magnetic torque, which increases as it approaches the helix. When the applied torque exceeds the resistive torque, the helix undergoes a screw-like motion, thereby enabling the stiffness tuning of the corresponding segment.

To determine the resistive torque threshold of the helical motion, the rotating magnet is moved toward the helix along the y-axis (Fig. 3a). Helical motion is initiated at a critical distance of 100 mm, corresponding to a field strength of ~0.5 mT (where T > T_m_resist). Maintaining a constant radial distance (y = 80 mm), we then translate the magnet laterally along the x-axis. Helical motion is observed when the lateral offset x_n decreases to 40 mm (n=1, 2, referring to the first and second magnetic helix).

Simultaneous or selective control of individual segments is achieved by adjusting the position of the actuation magnet. As shown in Fig. 3b, placing the magnet at x_1 = 5 mm and x_2 = 5 mm with a constant y (80 mm) triggers both helices, switching them to a high-stiffness state. Under x_1 = 36 mm and x_2 = 50 mm, the first segment is switched to a low-stiffness state. Likewise, repositioning x_2 to 31 mm (x_1 = 44 mm) reactivates the second helix, resulting in a decrease in the stiffness of the second segment.

By independently tuning the stiffness of the dual-segment module, the microcatheter tip can exhibit multiple bending modes under an S-field (20 mT), including C-shape (both segments in the medium-stiffness state), L-shape (first segment in the low-stiffness state and second segment in the high-stiffness state), and J-shape (first segment in the medium-stiffness state and second segment in the low-stiffness state) (Fig. 3c). Independent helix control enhances the adaptive dexterity of the Helixoft system in complex luminal structures.

Demonstration of the surgical capabilities of the Helixoft system in phantom lumens
To demonstrate its navigation and surgical functionality, the Helixoft-integrated microcatheter is controlled using a combined system comprising the aforementioned magnetic control module for steering and stiffness tuning, along with a self-developed robotic manipulator for precise advancement and withdrawal (Supplementary Fig. 7 and Supplementary Movie 3). The advancement speed is dynamically adjusted according to the luminal environment, with a speed range of 0–100 mm/s and a resolution of <0.5 mm. The parameters of the R-field and the S-field are modulated according to the theoretical model described above.

The navigation capability is first demonstrated using a 300 μm-diameter microcatheter to navigate a cerebrovascular phantom (ca. 1.5 mm in diameter) with successive bifurcations (Fig. 4a). The microcatheter initially advances smoothly in the straight main lumen but fails to enter the branch in its high-stiffness state. After being modulated to a low-stiffness state under the R-field, it successfully navigates into the left branch and reaches the secondary bifurcation under the S-field (20 mT). By changing the field direction, the microcatheter is then redirected into the right branch in the 2nd bifurcation, highlighting its adaptability to complex and branching vascular environments.

We further demonstrate the biopsy and payload delivery capabilities in a carotid phantom, featuring a phantom lesion for biopsy and a phantom aneurysm for payload delivery in the right and left branches, respectively. The phantom is fabricated from biocompatible hydrogel with a similar modulus (ca. 40 kPa) to real tissue, ideal for testing the interaction and safety of the Helixoft-integrated microcatheter58. As shown in Fig. 4b, the Helixoft-integrated microcatheter at the high-stiffness state advances smoothly without buckling in the straight section, controlled by the robotic manipulator. However, at the bifurcation, a 20-mT S-field fails to generate sufficient bending on the microcatheter due to its high stiffness, resulting in damage in the undesired position. To address this, the microcatheter is withdrawn and modulated to the low-stiffness state under the R-field. After that, the S-field is applied again, allowing the microcatheter to smoothly steer into either the right or left branch.

For biopsy, the Helixoft-integrated microcatheter is initially guided into the right branch at the low-stiffness state. However, the microcatheter buckles toward the phantom lesion (stained in red) and fails to puncture it due to insufficient rigidity (Fig. 4c). To address this, the microcatheter is switched back to the high-stiffness state and successfully penetrates the lesion in the end. This ability to freely switch among different stiffness states enables the Helixoft-integrated microcatheter to dynamically balance the flexibility for navigation and rigidity for mechanical intervention.

For payload delivery, the Helixoft-integrated microcatheter is magnetically controlled to perform multi-stage bending to enter the phantom aneurysm site (Fig. 4d). Initially, the microcatheter is steered with a primary bending of ca. 15° under the S-field to be aligned with the branch angle for smooth enter (Fig. 4b, t=29 s). As the tip approaches the bottom wall, a secondary bending of ca. 40° is introduced to prevent wall contact (t=30 s). This multi-stage bending allows the microcatheter to gradually adapt to the luminal environment, minimising the risk of tissue damage during navigation. Upon arrival at the target site, a mock drug (gelatin-based embolic agents with red ink) is injected through the fluidic tubing at microlitre-scale precision controlled by a microfluidic pump. This demonstration validates not only the precision drug delivery capability of the Helixoft system, but also its seamless integration and structural stability, particularly under bending stress.

Consequently, the Helixoft system demonstrates efficient navigation, localised biopsy and intraluminal payload delivery. Its scalable size, tuneable stiffness, active steering and non-disruptive control offer a versatile solution for surgical tasks in narrow and tortuous lumens with varying dimensions.

Validation of precision payload delivery in pulmonary interventional surgery of live pigs
Owing to its structural stability and compact size, the Helixoft-microcatheter is fully compatible with commercial real-time surgical imaging systems, such as X-ray fluoroscopy and endoscopy, and thus directly applicable in clinical procedures. In this section, we demonstrate the performance of precision payload delivery of the Helixoft-microcatheter and its quantitative safety evaluation during a classic pulmonary interventional procedure in live pigs (n=3). The microcatheter is magnetically controlled by a magnet embedded on a robotic arm with six revolute joints, and monitored in real-time by a digital radiography (DR) imaging setup (Fig. 5a). A commercial mini-camera is integrated with the Helixoft-microcatheter for the intratracheal visualisation of the bronchial anatomy during navigation (Fig. 5b).

The unique stiffness tuning capability of the Helixoft-integrated microcatheter enables efficient deep bronchial navigation, reaching the fourth-generation branches. Initially, the Helixoft-microcatheter is set to a high-stiffness state, enabling it to smoothly enter the trachea without buckling (t=0 s) (Fig. 5c). While this initial rigidity facilitates stable insertion in the trachea, it brings risks of mechanical irritation at bronchial bifurcations where the distal biopsy tip might scrape the lumen walls, potentially inducing bronchospasms. To mitigate this interference, the Helixoft-microcatheter is switched to the medium-stiffness state, increasing its compliance and thus minimising the mechanical mucosal irritation during navigation (t=15–45 s). The stiffness tuning process can be observed in the X-ray images of the pig in the prone position. Subsequently, magnetic steering is activated under the S-field to guide the microcatheter into the primary bronchus (t=60–75 s, lateral position). Continuing forward, it reaches another bifurcation and successfully enters the secondary bronchus (t=90–105 s). Owing to its small diameter, the microcatheter successfully accesses the tertiary bronchi and quaternary bronchi (ca. 2 mm in diameter, t=120–145 s). To validate the therapeutic delivery capability of the Helixoft system, an anticancer drug solution (Doxorubicin) containing contrast agent is administered through the fluidic tubing over a 15-s period (t=160–175 s) (Fig. 5d). We employ a fluidic control system to precisely regulate the amount of drug (Supplementary Fig. 8). A total of 25 μL of liquid drug is delivered through the microcatheter. The drug distribution is successfully monitored through both the integrated mini-camera and the X-ray radiography in the prone position (Supplementary Fig. 9). After administration, no residual drug remains in the catheter lumen, confirming the dosing accuracy. To assess the navigation success rate, ten independent navigation trials are performed in total. Under combined real-time X-ray and endoscopic imaging, all trials successfully reach the target position (Supplementary Figs. 10 and 11). The mean navigation time from the trachea to the quaternary bronchus is approximately 4 min.

The open architecture of our Helixoft system within compact size enables seamless integration with auxiliary dual-modality imaging modules, which allows real-time visualisation and precise spatio-temporal controllability of the microcatheter during multi-stage steering manoeuvres. Such synchronised control of a hybrid framework with imaging and actuation modules effectively prevents unintended contact with the lumen walls, ensuring smooth navigation, and minimising the risk of tissue damage. Post-operative assessment confirms the safety profile of the intervention, as indicated by maintenance of stable body weight and normal activity in the porcine models over a 7-day post-operative observation period (Supplementary Fig. 12). Furthermore, routine clinical blood analyses and histological examinations reveal no signs of acute inflammation or mucosal damage (Supplementary Fig. 13, Supplementary Note 3, and Supplementary Tables 3–6).

Demonstration of localised biopsy in real tissue
The Helixoft-integrated microcatheter combines programmable stiffness tuning with magnetic steering, enabling precise force interaction with delicate soft tissues. This capability is critical for performing localised biopsies within narrow and sensitive luminal environments, such as the oviduct, without damaging adjacent healthy tissue. In order to observe the real-time interaction details between the Helixoft system and physiological tissue, we use an ex-vivo porcine oviduct as the biopsy demonstration model (Fig. 6 and Supplementary Movie 4). This model provides a visually accessible and anatomically realistic biomechanical environment, with complex physiological features such as mucosal folds and mucus layers. Such a complex luminal environment introduces substantial friction and mechanical obstruction that impede the microcatheter’s navigation.

Traditional catheters often fail to maintain controllable tissue interaction during deep luminal biopsy missions, due to their inability to transition effectively between compliant steering and rigid penetration, frequently resulting in unintended buckling or tissue damage during tissue access or navigation. In contrast, our Helixoft-integrated microcatheter demonstrates adaptive performance through programmable stiffness control. In our demonstration, the Helixoft-integrated microcatheter uses its high-stiffness state to advance smoothly in the straight section of the oviduct until arriving at a corner (Fig. 6a). At this point, if no stiffness tuning is implemented, the magnetic tip collides with the lumen wall, causing unintended invasion of healthy tissue (t=2 s). To navigate the turn safely, the microcatheter is withdrawn and switched to the low-stiffness state under the R-field, allowing sufficient steering controlled magnetically under the S-field. After successfully navigating the lumen turn, the microcatheter is then switched to the medium-stiffness state, maintaining a ca. 25° bending angle to continue advancing. Upon reaching the target site, the microcatheter is pushed toward the target tissue but fails to penetrate due to buckling under insufficient stiffness. It is then switched back to the high-stiffness state. After that, it finally successfully penetrates the oviduct tissue (Fig. 6b). After biopsy, the microcatheter is withdrawn, carrying the obtained tissue sample in its biopsy needle. This biopsy demonstration highlights the programmable stiffness adaptability of the Helixoft-integrated microcatheter in complex physical environments, enabling safe and effective interaction with target sites while minimising disruption on environmental tissues.

Methods
Fabrication of the soft microtube
All chemicals used in this paper were purchased from Sigma-Aldrich (USA) unless otherwise specified. The soft microtube was fabricated using Ecoflex-20 (Dongzhixuan Co., Ltd., China), a biocompatible silicone-based elastomer. The soft microtube, with a thin wall thickness, was fabricated through a one-step moulding process on the commercial catheter. Specifically, two components of Ecoflex were mixed at a 1:1 weight ratio, and the Ecoflex mixture was transferred into a vacuum container to remove air bubbles. Afterwards, the Ecoflex mixture was filled between concentric microtube walls with defined diameter differences, allowing for a wide dimension range for the soft microtube (0.3–1 mm). Furthermore, the Ecoflex mixture in the tube was cured at 60 °C for 20 min. After moulding, the soft microtube was directly integrated onto the commercial microcatheter, with both ends of the soft microtube connected to the microcatheter tube. The inner and outer diameters of the soft microtube match those of the commercial microcatheter. Three dimensions of soft microtubes were fabricated, including 300 μm, 600 μm, and 1 mm.

Fabrication of the Helixoft-integrated microcatheter
After fabricating the soft microtube, a small anchor was glued onto the commercial microcatheter. The rigid magnetic helix was coaxially assembled with the Ecoflex tube, with the length of the helix being longer than that of the soft microtube. The soft microtube was connected to the commercial microcatheter (e.g., PE tube), which was fixed at the distal end. The rigid magnetic helix was fabricated by tightly winding a stainless-steel wire around a cylindrical metal rod, where the helix diameter was determined by the rod’s diameter. The pitch and number of turns were adjusted to tune the stiffness and length of the structure. The rigid magnetic helix, with higher stiffness, was coaxial with the soft microtube, which had lower stiffness. Three rigid magnetic microhelices with varying dimensions—0.5 mm (0.08-mm wire diameter, 0.3-mm pitch), 1.1 mm (0.2-mm wire diameter, 0.2-mm pitch), and 2 mm (0.1-mm wire diameter, 1-mm pitch)—were matched individually with corresponding soft microtubes. The tubular distal end was modular and material-selective, tailored to specific functional requirements. For tissue biopsy, a stainless-steel needle structure (outer diameter: 0.5 mm) was employed and partially filled with an NdFeB–Ecoflex composite (10 wt% NdFeB particles) to ensure magnetisation stability. The delivery channel for payloads was defined by a fibre inserted during the composite filling process, resulting in a channel approximately 100 μm in diameter. For applications without biopsy requirements, the distal head was fully enclosed within the microcatheter. All magnetic components were magnetised using a high-field magnetizer, with five magnetisation cycles to ensure long-term stability of magnetisation60,61. The magnetic components (stainless steel helix, biopsy needle with NdFeB–Ecoflex composite, ring-shaped NdFeB magnet, and pure biopsy needle) were characterised using vibrating sample magnetometry (Supplementary Fig. 14). A biocompatible lubricating layer of polydimethylacrylamide (PDMAA) was coated onto the surface of the microcatheter. The microcatheter was thoroughly cleaned by sequential rinsing with ethanol and isopropanol to remove surface contaminants. The cleaned microcatheter was treated with oxygen plasma (100 W, 1 min) to activate the surface, enhancing initiator attachment and polymer adhesion. The helix was pre-treated by silanisation using APTES (3-aminopropyltriethoxysilane). After this, the catheter was immediately immersed in an ethanol solution containing 10 wt% benzophenone for 20 min. This step allows benzophenone to adsorb onto the activated surface. The catheter was then transferred to an aqueous solution containing 20 wt% N,N′-dimethylacrylamide and 0.5 wt% Irgacure 2959. The microcatheter was soaked for 10 min to allow penetration of the monomer. Ultraviolet (UV) light (365 nm) was applied for 1 h to initiate in-situ polymerisation, resulting in the formation of a stable PDMAA coating on the catheter surface. The coated catheter was rinsed thoroughly with deionised (DI) water to remove any unreacted monomers or residual chemicals, then dried under nitrogen flow. The microcatheter was sterilised using ethylene oxide (EtO). This sterilisation process did not affect its magnetisation or structural integrity, as verified by surface magnetic flux intensity measurements, repeated bending tests, and payload delivery performance (Supplementary Fig. 15). Cytotoxicity and long-term adhesion of the hydrogel coating were evaluated by cell culture (BEAS-2B cell line, obtained from Saibaikang, Biological iCell-h023) and by bending-cycle tests in saline, respectively (Supplementary Note 4 and Supplementary Figs. 16 and 17).

Stiffness test
The stiffness test was based on the three-point bending testing method (MARK-10 EMS303). The force sensor had a force sensing range (0–2.5 N) and a resolution of 0.0005 N. In the experiment, the sensor’s motion speed was 3 mm/min, and the displacement was fixed. The different types of stiffness were realised by changing the overlapping length of the rigid magnetic helix and soft microtube. The stiffness states were defined by helix pitch coverage length (L₁=turns×pitch) together with the corresponding bending stiffness (EI), thereby avoiding ambiguity related to catheter diameter. Specifically, for the helix with a pitch of 1 mm and diameter of 2 mm, the three states were: Low stiffness: (pitch coverage L₁=0 mm, EI=2.06×10⁻⁸ N·m²), Medium stiffness: (pitch coverage L₁=3 mm, EI=3.60×10⁻⁸ N·m²), and High stiffness: (pitch coverage L₁=6 mm, EI=1.72×10⁻⁷ N·m²). For the helix with a pitch of 0.2 mm and diameter of 1 mm, the three states were: Low stiffness: (pitch coverage L₁=0 mm, EI=1.72×10⁻⁹ N·m²), Medium stiffness: (pitch coverage L₁=1.9 mm, EI=9.72×10⁻⁹ N·m²), and High stiffness: (pitch coverage L₁=3.8 mm, EI=7.43×10⁻⁸ N·m²). The control accuracy of the stiffness tuning and the microcatheter pose was furthermore analysed using a classic MAPE value test with MATLAB and Origin.

Magnetic control module
The magnetic control module employed either a coil system or a permanent magnet mounted on a robotic arm. The coil-based module consisted of three pairs of custom-built Helmholtz coils made of copper wire and connected to a direct current (DC) power supply. The three pairs of coils generated a three-axis magnetic field with 99% field homogeneity, producing field strengths up to 30 mT within a 50-mm spherical region over a frequency range of 0–20 Hz. The magnet-based module used a permanent block magnet (45×45×21 mm) mounted on a robotic arm, providing a maximum field strength of 200 mT. The robotic arm with six revolute joints was used to manoeuvre the block magnet. The magnetic control module generated the R- and S-fields for stiffness tuning and active steering. In the coil-based module, field strength and frequency were adjusted by varying the coil current, with parameters controlled through a graphical user interface. In the magnet-based module, the R- and S-fields were produced by rotating or reorienting the magnet with robotic arm control, and the R-field frequency was set by the motor rotation frequency (0–5 Hz). The coil-based module was mainly used for benchtop experiments and 3D phantom models, whereas the magnet-based module was applied in vivo to provide a larger working space. To accommodate different magnetic field penetration depths in vivo, magnets of varying sizes and field strengths were selected accordingly. The microcatheter’s bending was tested over 100 cycles, with a control accuracy of less than 1°. For the helix rotation, the strength of the R-field was 3 mT. For modelling, the strength of the S-field was set at 150 mT to test the maximum steering angle. The maximum steering angle was determined by gradually increasing the deflection angle until rebound of the magnetic tip was observed. For demonstrations in the phantom model and isolated organ, the strength of the S-field was maintained at 20 mT. For demonstration in the live pig, the strength of the S-field was 15 mT. The influence of the gradient field was weak below 20 mT (Supplementary Fig. 18).

Advancement based on the robotic catheter manipulator
A friction wheel mechanism was designed to drive the microcatheter. A pair of friction wheels was used to clamp the microcatheter. When the two friction wheels rotate in opposite directions simultaneously, forward or backward movement of the microcatheter is generated. The maximum driving speed of the friction wheels was 100 mm/s, with a positioning accuracy of better than 0.5 mm. Additionally, we designed a driver that adapts to the modular structural layout, achieving a high degree of integration for the wiring and sensors.

Payload delivery in the carotid phantom fabricated with tissue-mimicking hydrogel
The phantom was based on the solidified 10% gelatin. Firstly, the structure of the phantom with bifurcation and target area was designed by software. And a mould was printed using a 3D printer (Ultimaker S3, China). Then the 10% gelatin solution was filled into the mould. Then the mould was cooled at 4 °C. The gelatin-based embolic agent with red ink was chosen as the mock drug. The mock drug stored in a 1.5 mL tube can be pumped by a pressure pump (Fluigent). The flow speed can be controlled by adjusting the pressure. The pressure range was 0–2000 mPa. The sealing property of the microcatheter was tested under high pressure (2000 mPa) during the bending process. In addition, the microcatheter was compatible with a protective outer catheter (Supplementary Note 5 and Supplementary Fig. 19).

Animal experiment
Animal experiments were conducted in accordance with procedures approved by the Animal Experiment Ethics Committee (IACUC) of Shenzhen LingFu Top Biotech Co., Ltd (approval no. TOPGM-IACUC-2025-0012). Three six-month-old male Bama mini-pigs (about 30 kg) were purchased from Fuhao Experimental Animal Breeding Center (Beijing, China). The pigs were anesthetised with Zoletil 50 and maintained under isoflurane anesthesia. Prior to the procedure, the microcatheter was soaked in PBS and sterilised under ultraviolet light for one hour. The microcatheter, equipped with a mini-camera, was inserted into the pig’s bronchial tree through the tracheal cannula, and its movement within the bronchial tree was monitored using DR imaging. The rigid magnetic helix, controlled by an external magnetic field, enabled forward and backward motion to tune the microcatheter’s stiffness, facilitating navigation through bronchial bifurcations and entry into the target area for anticancer drug injection. The anticancer drug solution (Doxorubicin) containing contrast agent (barium sulphate) was used in this study. Upon completing the injection, the microcatheter was carefully withdrawn along the original path. Following the procedure, the pig was monitored for three days. Pathological analysis was conducted during the observation period, and the pig’s general condition and body weight were recorded throughout the treatment. A required amount of mock drug was pre-loaded into the reservoir connected to the microcatheter. The drug solution was then delivered through the microcatheter. Following delivery, an additional 100 μL of air (or saline)—corresponding to the microcatheter’s internal dead volume (V_d = π * (d_c/2)² * L_c, where d_c is the internal diameter and L_c is the length)—was injected as a flushing buffer to ensure complete payload delivery.

Peer review rounds
[{"round": 1, "reviewer_comments": {"reviewer_1": "The authors present a magnetically controlled mechanical distal end stiffness tuning mechanism for steerable magnetic catheters. The proposed design consists of a helical steel wire wrapped around a soft catheter tube and a radially magnetized magnet, also known as a needle, at the distal tip of the catheter. The distal end stiffness is tuned via the translational motion of the helical spring on the catheter using an external rotating magnetic field (R-field) generated by electromagnetic coils. The catheter steering is performed by an external permanent magnet attached to a 6 DoF robotic arm. The paper is well-written, and the figures and movies are informative. The results are reproducible. The literature on stiffness-tuning strategies is systematically described in the introduction. Especially the material- and structure-based distinction is beneficial for the interested audience. That being said, a couple of recent publications relevant to the proposed work are missing.\nThe proposed stiffness tuning method is sound and quite interesting to the field. It could provide new insights into magnetic robotic research. However, the manuscript requires additional details highlighting the real potential of the technique and further analysis in terms of magnetic actuation.\nI have a few comments on the claims about decoupled tuning, the limitation of the proposed design due to magnetization, the completeness of the magnetic actuation model, and the safety of the proposed design listed below:\n1. The stiffness tuning is not, in fact, decoupled from tip bending. The stiffness can only be tuned once the catheter is straight. Is it possible to demonstrate stiffness tuning while the catheter tip is already in a bent state, either due to external mechanical constraints or due to the S-field? A concentric tube pull-push-based distal end stiffness tuning method exists ( 10.1126/sciadv.adg6438 ), which can also be used while the distal end is in a bent state. What is the benefit of the proposed approach compared to the pull-push method?\n2. Could the proposed stiffness tuning method be used for tuning multiple catheters or multiple segments on a single catheter in a way that is similar to the (0.1126/sciadv.adq1951)? A multi-catheter and multiple-segment stiffness-tuning approach would add value to the proposed stiffness-tuning technique, differentiating it from the pull-push-based stiffness-tuning mentioned in the first comment.\n3. What is the material of the needle and helical spring? Medical-grade steel is generally a soft magnetic material. The magnetization easy axis of the helical spring should be closed to the radial plane. However, the needle should have an easy axis in the axial direction. 100 mT is strong enough to significantly change the magnetization direction of the needle, especially during the high stiffness mode of the proposed system. When magnetic torque is not high enough to bend the distal tip due to the high stiffness, the increased magnetic field would remagnetize the steel needle towards the axial direction. There is less risk of remagnetization for the spring because of the radial easy axis.\nThe authors must conduct a magnetization direction analysis considering the shape anisotropy direction under different S-field strengths and demonstrate that the applied S-field does not cause remagnetization of the needle or spring. If remagnetization is occurring, they should use a soft-magnetic magnetization model to calculate the magnetic torques (10.1109/TRO.2007.910775). For the demagnetization factor of the hollow needle, they can refer (10.1016/j.jmmm.2008.11.046).\nAlternatively, they can use a hard magnetic tip magnet with radial magnetization.\n4. The external permanent magnet would also apply a significant amount of magnetic gradient force. Authors should model the magnetic gradient of the magnet and include the contribution of the gradient force in their model.\n5. I am unsure how safe it is to navigate an exposed needle and a helical spring with exposed sharp edges through blood vessels. Could authors introduce an external thin catheter to increase safety? Could the proposed stiffness tuning mechanism also operate inside a thin catheter?\n6. There are minor typos—for instance, Helixfot.", "reviewer_2": "Major comments\nLimited conceptual novelty relative to prior art on magnetic, variable-stiffness catheters\nThe manuscript presents a helix-sliding sleeve that shortens/elongates a soft segment to achieve a 40-fold stiffness change, coupled with radial-magnet torque steering. Yet the key ingredients (i) a magnetically actuated rigid helix, (ii) sliding-sleeve stiffness modulation, and (iii) decoupled torque steering each appear in the literature individually or in combination. Representative examples include LMPA-based or SMP-based magnetic catheters with multi-segment stiffness switching (Chautems et al., 2020; Piskarev et al., 2023) as well as continuum robots that integrate ferromagnetic steering with mechanically programmable stiffness (Kim et al., 2019; Mao et al., 2024).\nThe manuscript, therefore, risks being seen as an incremental combination rather than a substantive advance.\n“40-fold continuous tuning” is insufficiently benchmarked\nFigure 2 reports 36 kPa → 1.5 MPa (≈42×) tuning on a 600 µm catheter, but no head-to-head comparison is made with state-of-the-art devices that already achieve ≥10–60× tuning with faster (<0.3 s) response (e.g., fibre-jamming catheters, LMPA phase-change devices). Without quantitative, side-by-side data, including transition time, energy budget, reversibility, and cyclic lifetime claims, superiority remains speculative.\nDecoupled control strategy not yet rigorously validated under realistic constraints\nThe authors rely on a radially magnetised helix for R-field translation and a radially magnetised distal tip for S-field bending (Fig. 2d–f). However, real anatomical navigation will expose the catheter to hysteretic friction and field-gradient perturbations from medical equipment. Demonstrations are limited to an open benchtop triaxial coil (≤30 mT) plus an external permanent magnet (≤150 mT). No experiments show: (i) simultaneous R- and S-field actuation; (ii) operation through conductive tissue layers; or (iii) motion inside MR/CT field environments, where eddy fields and safety regulations can severely restrict allowable dB/dt.\nIn-vivo validation is preliminary and incomplete\nOnly a single 30 kg pig is reported for bronchial drug delivery (Fig. 4), with no quantitative metrics on navigation success rate, dose accuracy, mucosal trauma, or acute inflammatory response.\nModelling assumptions may not capture critical nonlinearities\nThe constant-curvature beam model (Eq. 1) presumes a uniform EI and ignores the locally stiff/free transition that occurs as the helix overlaps the soft tube. This simplification explains the ≤10 % MAPE reported, but those tests were run in air with negligible distributed contact forces. In tortuous lumina, distributed wall contact introduces piecewise-loading conditions; without contact mechanics, the model’s predictive power for in-situ steering remains unproven.\nMinor comments\nTerminology – The term “Helixoft” is catchy but potentially confusing; please clarify in the Abstract that it is not a brand name but a lab prototype.\nTransition speed – The helix translates at ≈5 mm s⁻¹ (Fig. 2d); with a 10 mm soft segment, a full soft→stiff transition therefore requires ~2 s plus magnetic-arm repositioning time. Explicitly compare this with the <300 ms transition reported for fiber-jamming catheters.\nLubricious coating durability – Long-term adhesion and cytotoxicity of the PDMAA coating are not characterised. Provide peel/abrasion data over >100 bend cycles in saline at 37 °C.\nMagnetic safety – 150 mT static fields within 3–5 cm of the thoracic cavity warrant a discussion of pacemaker/ICD compatibility.\nSerilisation and re-use – The assembly includes stainless-steel helix + silicone tube glued to a PE catheter. How does ethylene-oxide or plasma sterilisation affect the glue joint and magnetic remanence?\nNomenclature of stiffness states – “Low, medium, high” states correspond to specific helix rotations (0, 3, 6 turns), but the absolute EI values vary with catheter diameter. Report both the helix pitch coverage and the resulting EI to avoid ambiguity.\nOverall assessment\nThe manuscript and its conceptual advance appear modest when set against the rapidly moving field of magnetically actuated, variable-stiffness continuum robots. To reach the novelty and depth expected for Nature Communications, the authors should:\n1. Demonstrate a capability that prior devices cannot match—e.g., real-time decoupling under simultaneous R- and S-field actuation inside tissue-mimicking phantoms with distributed constraints, or a statistically significant performance gain (>2×) in task-completion time over a leading benchmark catheter.\n2. Provide a rigorous, quantitative benchmarking table that covers stiffness range, transition time, bending angle per Tesla, force throughput, and device OD across at least five state-of-the-art systems.\n3. Expand in-vivo validation to at least n = 3–5 large-animal procedures with histological assessment and device-specific adverse-event reporting.\nAddressing these points would substantially strengthen the manuscript and clarify its unique contribution beyond what is already available in the literature.", "reviewer_3": "This paper describes a microrobotic system that uses a helix-shaped magnetic soft microrobot positioned around a commercially catheter. Soft microrobot enabling continuous stiffness tuning and precise steering. The approach is interesting however there are some inconsistences. That I will describe in detail as follow\n1. Figure 1 is a diagram of a person inside of a 3-coil arrangement. The validation of the system was performed in an ex-vivo oviduct biopsy and in-vivo drug delivery to the bronchi of a live pig. In this way the scheme in figure 1 is not consistent with the experiments.\n2. To generate the magnetic actuation in this paper they use a 3 coils system and a permanent magnet mounted on a robotic ar. Usually the rotating magnetic field is generated by a Helmholtz coil array consisting of three pairs of coils or something similar. This can of system con perfectly control the direction of magnetic field, why they use a permanent magnet mounted on a robotic arm.A detailed description of the design and operation of the magnetic actuation systems should be included in the Materials and Methods section, as well as photographs in the Supporting Information.\n3. Importantly, the paper lacks crucial information regarding the experiments. For example, there's no detailed description of the magnetic controller and its functionality, nor is the material composition and fabrication method of the rigid helical microrobot. Seems that they use stainless steel with magnetic properties… The authors need to include this information in the materials and methods section\n4. The paper fails to describe the Helixoft microrobot's morphology characterization. Crucially, it omits information regarding the composition of its magnetic components (percentage and type) and the magnetization curve of its ferromagnetic material.\nOverall, the research contains a lot of experimental data, but it's poorly organized and shouldn't be published in its current state."}, "author_response": "Comment 1: The authors present a magnetically controlled mechanical distal end stiffness tuning mechanism for steerable magnetic catheters. The proposed design consists of a helical steel wire wrapped around a soft catheter tube and a radially magnetized magnet, also known as a needle, at the distal tip of the catheter. The distal end stiffness is tuned via the translational motion of the helical spring on the catheter using an external rotating magnetic field (R-field) generated by electromagnetic coils. The catheter steering is performed by an external permanent magnet attached to a 6 DoF robotic arm.\nThe paper is well-written, and the figures and movies are informative. The results are reproducible. The literature on stiffness-tuning strategies is systematically described in the introduction. Especially the material- and structure-based distinction is beneficial for the interested audience. That being said, a couple of recent publications relevant to the proposed work are missing.\nResponse to comment 1：\nWe thank the reviewer for the positive feedback and recognition of our work, as well as the suggestion regarding relevant recent publications. In response, a more comprehensive literature search was conducted, and the manuscript has been updated to include several recent studies from the past two years on stiffness modulation mechanisms and magnetic control strategies. The relevant references have been added to the revised manuscript to better support the context of our approach.\nThe following references have been newly added to the manuscript (Pages 32 and 33):\n“57.Wang, C. et al. Heterogeneous multiple soft millirobots in three-dimensional lumens. Sci. Adv. 10, eadq1951 (2024).\n60.Abbott, J. J., Ergeneman, O., Kummer, M. P., Hirt, A. M. & Nelson, B. J. Modeling magnetic torque and force for controlled manipulation of soft-magnetic bodies. IEEE Trans. Robot. 23, 1247–1252 (2007).\n61.Beleggia, M., Vokoun, D. & De Graef, M. Demagnetization factors for cylindrical shells and related shapes. J. Magn. Magn. Mater. 321, 1306–1315 (2009).”\nComment 2: The proposed stiffness tuning method is sound and quite interesting to the field. It could provide new insights into magnetic robotic research. However, the manuscript requires additional details highlighting the real potential of the technique and further analysis in terms of magnetic actuation.\nI have a few comments on the claims about decoupled tuning, the limitation of the proposed design due to magnetization, the completeness of the magnetic actuation model, and the safety of the proposed design listed below:\n1. The stiffness tuning is not, in fact, decoupled from tip bending. The stiffness can only be tuned once the catheter is straight. Is it possible to demonstrate stiffness tuning while the catheter tip is already in a bent state, either due to external mechanical constraints or due to the S-field? A concentric tube pull-push-based distal end stiffness tuning method exists ( 10.1126/sciadv.adg6438 ), which can also be used while the distal end is in a bent state. What is the benefit of the proposed approach compared to the pull-push method?\nResponse to comment 2:\nWe thank the reviewer for the insightful comment concerning stiffness tuning of the microcatheter tip in a bent state. In response, we conducted a new experiment to demonstrate microcatheter stiffness tuning under this condition.\nAs shown in Fig. R1, the microcatheter initially bends in a low-stiffness state. When exposed to an R-field, the rigid helix undergoes a helical motion, gradually overlapping with the soft microtube and thereby tuning the overall stiffness of the microcatheter. However, stiffness tuning during bending may induce additional deflection and generate localised contact forces, thereby increasing the risk of tissue damage during navigation. Therefore, such simultaneous tuning is not the focus of our approach. For clinical safety, the system is designed to operate in a sequential manner across distinct modes, including stiffness tuning, advancement, steering, and biopsy. Before each operation, the microcatheter stiffness is tuned under the R-field to the appropriate state. During advancement through straight luminal segments, the catheter is set to a high-stiffness state to prevent distal buckling. When approaching bifurcations or traversing tortuous lumens, its stiffness is reduced under the R-field so that the magnetic distal end can bend the soft microtube under the S-field for active steering. Upon reaching the target site, the microcatheter is switched back to a high-stiffness state to provide sufficient mechanical stability for biopsy or other manipulations.\nFig. R1 Screw-like helical motion of the helix with the microcatheter tip maintained in a bent state.\nScale bar: 5 mm.\nWe note that the reference (10.1126/sciadv.adg6438)1 reports magnetic navigation of a guidewire within a 7 T MRI environment. This work has been cited in the revised manuscript as an important reference for magnetic navigation strategy (ref. 59). However, this study offers only a limited discussion of stiffness modulation.\nThe pull–push method is based on concentric tubes, as reported in the reference (10.1109/TRO.2024.3378442)2. Device miniaturisation remains challenging due to the assembly of proximally controlled multilayer continuum tubes. Moreover, inserting a stiffer inner tube into a bent soft tube inevitably induces additional deflection3, which may generate undesired forces on the surrounding tissue, especially in narrow lumens. Compared with conventional pull-push concentric tube designs, our Helixoft-integrated system is downsized to 300 μm in diameter while maintaining high compliance with precise, continuous stiffness tuning to minimise tissue trauma. Besides, it achieves decoupling of stiffness tuning and active steering using a single magnetic field with remote controllability and intrinsic safety.\nWe have added a comparative discussion on conventional variable-stiffness catheters in the Introduction section (Page 3):\n“Moreover, previous catheters based on material- and structure-based strategies were primarily developed with diameters exceeding 2 mm, a compromise size requirement to accommodate bulky actuation components, which in turn restricts their applicability mainly to cardiac and gastrointestinal interventions. To date, downsizing stiffness-variable catheters to diameters approaching 1 mm—a critical threshold for vascular and reproductive system applications—remains a significant challenge3,28,53.”\nWe have also included a description of the control strategy underlying stiffness tuning and active steering in the Results section (Page 6):\n“For clinical safety, the system is designed to operate in a sequential manner across distinct modes, including stiffness tuning, advancement, steering, and biopsy. Before each operation, the microcatheter stiffness is tuned under the R-field to the appropriate state. During advancement through straight luminal segments, the catheter is set to a high-stiffness state to prevent distal buckling. When approaching bifurcations or traversing tortuous lumens, its stiffness is reduced under the R-field so that the magnetic distal end can bend the soft microtube under the S-field for active steering. Upon reaching the target site, the microcatheter is switched back to a high-stiffness state to provide sufficient mechanical stability for biopsy or other manipulations.”\nComment 3: 2. Could the proposed stiffness tuning method be used for tuning multiple catheters or multiple segments on a single catheter in a way that is similar to the (0.1126/sciadv.adq1951)? A multi-catheter and multiple-segment stiffness-tuning approach would add value to the proposed stiffness-tuning technique, differentiating it from the pull-push-based stiffness-tuning mentioned in the first comment.\nResponse to Comment 3:\nThanks for the reviewer’s valuable suggestion on the multi-catheter and multiple-segment stiffness-tuning approach. Inspired by the reviewer’s suggestion, we demonstrated the scalability of our Helixoft system by implementing a multi-segment design that enables multi-level stiffness tuning across different regions of a single microcatheter, adopting a control strategy similar to that reported in (10.1126/sciadv.adq1951)4. It has also been cited as an important reference supporting the independent control of multiple components (ref. 57). As shown in Fig. R2, a dual-segment stiffness-tuning module is integrated along the microcatheter. Each segment consists of a rigid magnetic helix coupled to a soft microtube. Independent control of each helix is achieved via the interplay between the magnetic actuation and local friction. A permanent cylindrical magnet (1.5 cm in diameter) generates a rotating magnetic torque, which increases as it approaches the helix. When the applied torque exceeds the resistive torque, the helix undergoes a screw-like motion, thereby enabling the stiffness tuning of the corresponding segment.\nTo determine the resistive torque threshold of the helical motion, the rotating magnet is moved toward the helix along the y-axis (Fig. R2a). Helical motion is initiated at a critical distance of 100 mm, corresponding to a field strength of ~0.5 mT (where T T ). Maintaining a constant radial distance (y=80 mm), we then translate the magnet laterally along the x-axis. Helical motion is observed when the lateral offset x decreases to 40 mm (n = 1.2, referring to the first and second magnetic helix).\nn\nSimultaneous or selective control of individual segments is achieved by adjusting the position of the actuation magnet. As shown in Fig. R2b, placing the magnet at x 1 =5 mm and x =5 mm with a constant y (80 mm) triggers both helices, switching them to a high-stiffness state. Under x =36 mm and x =50 mm, the first segment 1 2 is switched to a low-stiffness state. Likewise, repositioning x to 31 mm ( x 2 1 =44 mm) reactivates the second helix, resulting in a decrease in the stiffness of the second segment.\nFig. R2 Distributed stiffness control with a dual-segment stiffness-tuning module. a, Strategy for estimating resistive torque to enable independent helix control. Error bars represent standard deviation, n=3 times, mean±sd. b, Independent control of two helices in a single microcatheter. c, Multiple bending configurations of the corresponding microcatheter. Scale bar: 5 mm.\nBy independently tuning the stiffness of the dual-segment module, the microcatheter tip can exhibit multiple bending modes under an S-field (20 mT), including C-shape (both segments in the medium-stiffness state), L-shape (first segment in the low-stiffness state and second segment in the high-stiffness state), and J-shape (first segment in the medium-stiffness state and second segment in the low-stiffness state) (Fig. R2c). Independent helix control enhances the adaptive dexterity of the Helixoft system in complex luminal structures.\nThe design of the dual-segment stiffness-tuning configuration has been added as a new section in the manuscript (Pages 10-12), as follows:\n“Distributed stiffness control with a dual-segment module\nTo demonstrate the scalability of our stiffness-tuning mechanism, we have implemented a multi-segment design that enables multi-level stiffness tuning across different regions of a single microcatheter57. As shown in Fig. 3, a dual-segment stiffness-tuning module is integrated along the microcatheter. Each segment consists of a rigid magnetic helix coupled to a soft microtube. Independent control of each helix is achieved via the interplay between the magnetic actuation and local friction. A permanent cylindrical magnet (1.5 cm in diameter) generates a rotating magnetic torque, which increases as it approaches the helix. When the applied torque exceeds the resistive torque, the helix undergoes a screw-like motion, thereby enabling the stiffness tuning of the corresponding segment.\nTo determine the resistive torque threshold of the helical motion, the rotating magnet is moved toward the helix along the y-axis (Fig. 3a). Helical motion is initiated at a critical distance of 100 mm, corresponding to a field strength of ~0.5 mT (where T T ). Maintaining a constant radial distance (y=80 mm), we then translate the magnet laterally along the x-axis. Helical motion is observed when the lateral offset x decreases to 40 mm (n = 1.2, referring to the first and second magnetic helix).\nn\nSimultaneous or selective control of individual segments is achieved by adjusting the position of the actuation magnet. As shown in Fig. 3b, placing the magnet at x 1 =5 mm and x =5 mm with a constant y (80 mm) triggers both helices, switching them to a high-stiffness state. Under x =36 mm and x =50 mm, the first segment 1 2 is switched to a low-stiffness state. Likewise, repositioning x to 31 mm ( x 2 1 =44 mm) reactivates the second helix, resulting in a decrease in the stiffness of the second segment.\nBy independently tuning the stiffness of the dual-segment module, the microcatheter tip can exhibit multiple bending modes under an S-field (20 mT), including C-shape (both segments in the medium-stiffness state), L-shape (first segment in the low-stiffness state and second segment in the high-stiffness state), and J-shape (first segment in the medium-stiffness state and second segment in the low-stiffness state) (Fig. 3c). Independent helix control enhances the adaptive dexterity of the Helixoft system in complex luminal structures.”\nComment 4: 3. What is the material of the needle and helical spring? Medical-grade steel is generally a soft magnetic material. The magnetization easy axis of the helical spring should be closed to the radial plane. However, the needle should have an easy axis in the axial direction. 100 mT is strong enough to significantly change the magnetization direction of the needle, especially during the high stiffness mode of the proposed system. When magnetic torque is not high enough to bend the distal tip due to the high stiffness, the increased magnetic field would remagnetize the steel needle towards the axial direction. There is less risk of remagnetization for the spring because of the radial easy axis.\nThe authors must conduct a magnetization direction analysis considering the shape anisotropy direction under different S-field strengths and demonstrate that the applied S-field does not cause remagnetization of the needle or spring. If remagnetization is occurring, they should use a soft-magnetic magnetization model to calculate the magnetic torques (10.1109/TRO.2007.910775). For the demagnetization factor of the hollow needle, they can refer (10.1016/j.jmmm.2008.11.046).\nAlternatively, they can use a hard magnetic tip magnet with radial magnetization.\nResponse to comment 4:\nWe thank the reviewer for the insightful comments regarding the magnetisation stability of the needle and helical spring under S-fields. In our design, the biopsy needle is fabricated from medical-grade stainless steel, while the helix is made of spring steel. Magnetisation stability of the needle is crucial for preventing unintended bending during helical motion. The magnetic components are magnetised under a 1.1 T field. As shown in Fig. R3, the coercive field of the pure biopsy needle is approximately 60 mT, which is sufficient to meet the requirements of most clinical scenarios and the application conditions of our experiments. To ensure stable axial magnetisation, a thin NdFeB–Ecoflex composite layer is further coated on the inner wall of the needle lumen, increasing the coercive field to 600 mT.\nFig. R3 Magnetisation curves of the magnetic components.\nThe references (10.1109/TRO.2007.9107755 and 10.1016/j.jmmm.2008.11.0466) suggested by the reviewer are highly relevant to the application of magnetic catheters. As the reviewer’s suggestion, we updated the modelling of the magnetic torque under the remagnetization conditions in Supplementary Information (Pages 9), as follows:\n“If remagnetisation occurs, the general form of the magnetic torque can be expressed as13,14:\nT VMB (11)\nWhen the applied field is below the saturation threshold of the biopsy needle material, magnetisation increases approximately linearly with field strength. In this low-field regime, the magnetic torque therefore scales quadratically with the applied field:\nV D D\nT  r a B 2 sin（2） (12)\n2D D\n0 a r\nWhen the applied field is strong enough to bring the magnetisation into saturation, the torque becomes independent of further increases in field strength.\nV D D\nT  r a m 2sin(2) (13)\n2 s\nB\n（D -D）m sin(2)2 sin() (14)”\nr a s \n“The rigid magnetic helix was fabricated by tightly winding a stainless-steel wire around a cylindrical metal rod, where the helix diameter was determined by the rod’s diameter. The pitch and number of turns were adjusted to tune the stiffness and length of the structure. The rigid magnetic helix, with higher stiffness, was coaxial with the soft microtube, which had lower stiffness. Three rigid magnetic microhelices with varying dimensions—0.5 mm (0.08-mm wire diameter, 0.3-mm pitch), 1.1 mm (0.2-mm wire diameter, 0.2-mm pitch), and 2 mm (0.1-mm wire diameter, 1-mm pitch)—were matched individually with corresponding soft microtubes. The tubular distal end was modular and material-selective, tailored to specific functional requirements. For tissue biopsy, a stainless steel needle structure (outer diameter: 0.5 mm) was employed and partially filled with an NdFeB–Ecoflex composite (10 wt% NdFeB particles) to ensure magnetisation stability. The delivery channel for payloads was defined by a fibre inserted during the composite filling process, resulting in a channel approximately 100 μm in diameter. For applications without biopsy requirements, the distal head was fully enclosed within the microcatheter. All magnetic components were magnetised using a high-field magnetizer, with five magnetisation cycles to ensure long-term stability of magnetisation60,61. The magnetic components (stainless steel helix, biopsy needle with NdFeB–Ecoflex composite, ring-shaped NdFeB magnet, and pure biopsy needle) were characterised using vibrating sample magnetometry (VSM) (Supplementary Fig. 14).”\nComment 5: 4. The external permanent magnet would also apply a significant amount of magnetic gradient force. Authors should model the magnetic gradient of the magnet and include the contribution of the gradient force in their model.\nResponse to comment 5:\nThank you for the reviewer’s insightful suggestion regarding the influence of magnetic gradient forces on the modelling. A permanent magnet can induce a magnetic gradient force. For helical motion, a cylindrical anchor is affixed to the outer layer of the soft microtube, positioned between the threads of the rigid magnetic helix. This anchoring structure prevents undesired translational slipping of the rigid helix caused by gradient forces.\nDuring the bending process, the magnetic field gradient can induce translational motion of the magnetic tip. The magnetic gradient force acting on a biopsy needle can be expressed as7,8:\n B B \ncos 2x sin 2x\n \nx y\nF (M )B M   (1)\ngrad 2  B B \ncos 2y sin 2y \n x y \nThe normal component of this gradient force that contributes to bending is:\n B B B B \nF F nM  2x cossin 2x sin2 2y cos2 2y cossin\n \nn grad  x y x x \n(2)\nsin\nwhere n is the normal vector, M is the magnetic moment of the biopsy\n \ncos \nneedle, and B is the magnetic gradient of the S-field.\nM  F L (3)\ngrad n\nHere, L is the lever arm length from the point of force application to the bending axis.\nThe equilibrium condition balancing the magnetic torque, gradient-induced torque, and the bending torque, is given by:\n\nMB sin（r-）+F L EI (4)\n2 n L\nAs the distance from the external magnet increases, the rapid decay of the magnetic field gradient diminishes the gradient force acting on the distal tip, especially in deep tissue regions. Under such conditions, the bending induced by the gradient force is effectively negligible. To evaluate the contribution of the gradient force to bending, we performed a finite element analysis (FEA) as shown in Fig. R4a. The magnetic field strength at the tip was kept constant at 20 mT and 30 mT to examine the effect of field gradients. By comparing the resulting bending angles under these conditions, the influence of the gradient force was assessed. As shown in Figs. R4b and R4c, the bending is only minimally affected when the field strength falls below 20 mT.\nFig. R4 Comparison of bending angle under the uniform field and the gradient field. a, Finite element analysis (FEA) of bending under a uniform magnetic field versus a gradient field. b, Bending simulation results under 30 mT and 20 mT magnetic fields. c, Simulated bending results under varying S-field directions at a constant 20 mT field strength.\nThe influence of the gradient field on the modelling is added in Supplementary Fig. 18, and the discussion has been updated in the Supplementary Information (Page 11):\n“For the field generated by the permanent magnet, the bending is only minimally affected when the field strength falls below 20 mT, as illustrated in Supplementary Fig. 18.”\nComment 6: 5. I am unsure how safe it is to navigate an exposed needle and a helical spring with exposed sharp edges through blood vessels. Could authors introduce an external thin catheter to increase safety? Could the proposed stiffness tuning mechanism also operate inside a thin catheter?\nResponse to comment 6:\nWe thank the reviewer for this insightful comment regarding the safety of an exposed needle and helix within blood vessels. In vascular applications with only payload delivery requirements, a magnetic tube rather than a biopsy needle is adopted (Fig. R5). Besides, the magnetic tube can be sheathed within the microcatheter and not directly exposed in the blood vessel lumen. We consider that the reviewer’s suggestion of introducing an external thin catheter is very valuable, as adding a protective catheter can also enhance the safety during pulmonary interventional surgery.\nThus, we implemented an additional concentric PTFE catheter with a wall thickness of 100 μm, serving as a protective layer of the Helix-integrated system (Fig. R5). The entire assembly can advance as a single unit. The Helixoft system is pushed out of the protective catheter to deliver payloads and then be withdrawn. Furthermore, the microcatheter system can be switched to a low stiffness state within the protective catheter under an R-field. The microcatheter is extended beyond the protective catheter and performs active steering under an S-field. Thus, the Helixoft system with a thin protective catheter can retain its functionality, including stiffness tuning, active steering, and efficient payload delivery.\nFig.R5 Operation sequence and functionality of the Helixoft-integrated microcatheter with a concentric protective catheter. The assembly, consisting of a microcatheter sheathed within a 100 μm-thick outer catheter, advances as a unit. The microcatheter—initially in a stiffened state—is extended beyond the protective catheter for precise payload delivery and subsequently retracted. For navigation, the microcatheter is softened while inside the protective catheter, then deployed and magnetically steered. The protective catheter enhances safety while preserving the key functions of the microcatheter, including continuous stiffness tuning, active steering, and payload delivery. Scale bar: 5 mm.\nWe have added the compatibility with a protective outer catheter in Supplementary Fig. 19 of the Supplementary Information. The relevant content has been added in Supplementary Note 5 of the Supplementary Information (Page 13):\n“Supplementary Note 5 | Compatibility with a protective outer catheter\nOwing to the high scalability of our system, we implement an additional concentric PTFE catheter with a wall thickness of 100 μm, which serves as a protective layer for the Helixoft-integrated system (Supplementary Fig. 19). The entire assembly advances as a single unit. The Helixoft system is pushed out of the protective catheter to deliver payloads and then be withdrawn. Moreover, the microcatheter system can be switched to a low-stiffness state within the protective catheter under an R-field. The microcatheter extends beyond the protective catheter and performs active steering under an S-field. Thus, the Helixoft system with a thin protective catheter can perform the required functions, including stiffness tuning, active steering, and efficient payload delivery.”\nComment 7: 6. There are minor typos—for instance, Helixfot.\nResponse to comment 7:\nThanks for the valuable comment on the typos. We have carefully reviewed the entire manuscript and have revised typos on Page 14, as follows:\nPage 15: “Attribute to” has been revised as “Owing to”.\nPage 16: “Helixfot” has been replaced by“Helixoft”.\nPage 22: “salinisation” has been revised as “silanisation”.\nPage 23: “type” has been revised as “types”.\nReviewer #2\nComment 1: “40-fold continuous tuning” is insufficiently benchmarked\nFigure 2 reports 36 kPa → 1.5 MPa (≈42×) tuning on a 600 µm catheter, but no head-to-head comparison is made with state-of-the-art devices that already achieve ≥10–60× tuning with faster (<0.3 s) response (e.g., fibre-jamming catheters, LMPA phase-change devices). Without quantitative, side-by-side data, including transition time, energy budget, reversibility, and cyclic lifetime claims, superiority remains speculative.\nResponse to comment 1:\nWe thank the reviewer for pointing out the importance of benchmarking our variable-stiffness performance against existing state-of-the-art variable-stiffness catheter technologies. In response, we prepared a detailed table summarising the research framework of variable-stiffness and steerable catheters, as also suggested in Comment 12. Our work is based on valuable insights gained from previous studies. While these earlier studies each offer distinct advantages, they also face significant challenges, particularly regarding device outer diameter, safety, and remote controllability. For example, fibre-based catheters can achieve rapid stiffness modulation, but their miniaturisation is limited to ~2.3 mm in diameter9. LMPA-based systems can provide a broader stiffness range (≥10–60×)10. However, their reliance on intense thermal or electrical stimulation raises safety concerns in sensitive luminal environments, including risks of material toxicity, thermal injury, and electrical damage. Moreover, the reported minimum dimensions of such devices remain at ~1 mm, with the constraints of their actuation mechanisms hindering further miniaturisation and integration at smaller scales (Table R1).\nOur system is specifically designed for narrow and sensitive lumens (e.g., bronchioles, cerebral vasculature, and fallopian tube) through downsizing the variable-stiffness catheter to 300 μm in diameter. Independent remote magnetic control of stiffness tuning and active steering is achieved through a decoupling strategy that combines helical motion with torque-driven bending of individual microrobot components. This approach avoids other potentially harmful stimuli, such as thermal, optical, or electrical inputs. Moreover, repeated in vivo validation across different animal models and systematic safety evaluations confirm the feasibility and safety of the Helixoft system. These key advantages have been included in Table R1.\nFor transition time and reversibility, we quantify the durations required for softening and stiffening (Table R1). Regarding the energy budget and cyclic lifetime, a quantitative comparison with other approaches (e.g., pneumatic or thermal actuation) is inherently challenging due to the fundamentally different stiffness-modulation mechanisms. As our system is still at the research and validation stage, these aspects have not been a primary focus, but we acknowledge their importance and will investigate them systematically in future work. It is also worth noting that existing methods that rely on Joule heating inherently involve thermal losses, which may cause tissue injury in vivo.\nThe comparison table has been added as Supplementary Table R2. In addition, the following sentence has been inserted into the Introduction section (page 4):\n“Supplementary Table 2 summarises the performance of the Helixoft-integrated microcatheter in comparison with representative variable-stiffness and steerable catheters.”\nTable R1. Comparison of the Helixoft-integrated microcatheter with existing variable-stiffness and steerable catheters.\nDevice type Device Safe and Continuous In vivo Transition time stiffness Bending Force\nouter remote stiffness validation Softening Stiffening r ange angle per throughput\ndiameter control & safety Tesla\n≤1 mm\nMagnetically controlled 300-1000 Yes Continuous Yes(3 cases 0.3 s 0.3 s 40-fold 787 1.5 N\nhelical motion (This work) μm (0.12 kPa/dewith safety (118°/150\ng) assessment) mT)\nThermal-responsive low- 1 mm No Stepwise No 16 s 30 s 400-fold N/A 8 N\nmelting-point alloy10\n1-3 mm in device diameter\nVacuum-responsive fibre 2.3 mm No Stepwise No 0.1 s 0.3 s 6.5-fold 1337(107°N/A\njamming9 /80 mT)\nThermal-responsive low- 2.8/5.7 No Stepwise No N/A N/A N/A N/A N/A\nmelting-point alloy11 mm\nThermal-responsive low- 2.33 mm No Stepwise No 15 s 80 s N/A N/A N/A\nmelting-point alloy12\nThermal-responsive shape 2.3 mm No Stepwise No 13 s 4.4 s 66-fold 1587(127°N/A\nmemory polymer13 /80 mT)\nThermal-responsive 2 mm No Stepwise No 175 s 107 s 21-fold 637(51°/8 N/A\nconductive shape memory 0 mT)\npolymer14\nConcentric tube assembly15 1.8 mm No Continuous No N/A N/A 7.2-fold N/A N/A\n3-5 mm in device diameter\nThermal-responsive low- ~3-4 mm No Stepwise Yes(1 case 10 s 15 s N/A N/A N/A\nmelting-point alloy16 without\nsafety\nassessment)\nConcentric tube assembly2 3.5 mm No Stepwise No N/A N/A 14-fold N/A ~2.5 N\nThermoset shape memory 3 mm No quasi- No 9 s 60 s 93-fold 2250(90°/ N/A\npolymer17 continuous 80 mT)\n>5 mm in device diameter\nConcentric tube assembly18 7 mm No Stepwise No N/A N/A N/A 6750(135°N/A\n/20 mT\nNotes: Compared with existing variable-stiffness and steerable catheters, our Helixoft-integrated system is downsized to 300 μm and enables continuous stiffness tuning, making it suitable for narrow and sensitive lumens such as the bronchioles, cerebral vasculature, and fallopian tube. Moreover, the system achieves decoupling between stiffness tuning and active steering, both realised under a single magnetic field without introducing any additional potentially harmful stimuli.\nComment 2: Decoupled control strategy not yet rigorously validated under realistic constraints\nThe authors rely on a radially magnetised helix for R-field translation and a radially magnetised distal tip for S-field bending (Fig. 2d–f). However, real anatomical navigation will expose the catheter to hysteretic friction and field-gradient perturbations from medical equipment. Demonstrations are limited to an open benchtop triaxial coil (≤30 mT) plus an external permanent magnet (≤150 mT). No experiments show: (i) simultaneous R- and S-field actuation; (ii) operation through conductive tissue layers; or (iii) motion inside MR/CT field environments, where eddy fields and safety regulations can severely restrict allowable dB/dt.\nResponse to Comment 2:\nWe sincerely thank the reviewer for this important comment regarding the influence of the realistic constraints. In response, we implemented the following discussions:\nInfluence of friction hysteresis:\nDuring navigation, the microcatheter experiences tangential friction from the lumen wall, which primarily resists axial advancement but has only a limited effect on active steering. Under the imaging feedback, the magnetic field precisely controls the Helixoft system to dynamically modulate both its position and bending angle to adapt to the lumen environment and minimise excessive tip–wall contact. In addition, the catheter surface is coated with a low-friction hydrogel layer, minimising the mechanical resistance during navigation. Validation in hydrogel phantoms, ex vivo tissue, and live porcine models has confirmed the Helixoft’s high-performance navigation in complex lumens.\nThe discussion of friction hysteresis has been in the Discussion section on Page 20, as follows:\n“During navigation, the microcatheter also experiences tangential friction from the lumen wall, which primarily resists axial advancement but has only a limited effect on active steering. Under the imaging feedback, the magnetic field precisely controls the Helixoft system to dynamically modulate both its position and bending angle to adapt to the lumen environment and minimise excessive tip–wall contact. In addition, the catheter surface is coated with a low-friction hydrogel layer, minimizing the mechanical resistance during navigation. Validation in hydrogel phantoms, ex vivo tissue, and live porcine models have confirmed the Helixoft’s high-performance navigation in complex lumens.”\nSimultaneous R- and S-field actuation:\nR- and S-fields are employed to achieve stiffness tuning and active steering, respectively. The Helixoft microrobot is designed to decouple these functions, prioritising procedural safety, stability, and efficiency over simultaneous actuation of both fields. Although superposition of the two fields is theoretically feasible, as also suggested in Comment 11. For example, we have demonstrated the stiffness tuning of the microcatheter tip in a bent state (Fig. R1). However, we have not identified a specific clinical scenario that would require such an operation. Stiffness tuning and active steering are executed according to a well-defined temporal sequence, which ensures microcatheter stability and minimises tissue risk. In clinical practice, the Helixoft-integrated microcatheter remains soft during steering under the S-field and switches to a stiff state during advancement and biopsy under the R-field.\nWe have added the control strategy on Page 6 of the revised manuscript, as follows:\n“For clinical safety, the system is designed to operate in a sequential manner across distinct modes, including stiffness tuning, advancement, steering, and biopsy. Before each operation, the microcatheter stiffness is tuned under the R-field to the appropriate state. During advancement through straight luminal segments, the catheter is set to a high-stiffness state to prevent distal buckling. When approaching bifurcations or traversing tortuous lumens, its stiffness is reduced under the R-field so that the magnetic distal end can bend the soft microtube under the S-field for active steering. Upon reaching the target site, the microcatheter is switched back to a high-stiffness state to provide sufficient mechanical stability for biopsy or other manipulations.”\nOperation through conductive tissue layers:\nMagnetic fields exhibit high penetrability through biological tissues, including conductive layers19,20. Unlike electric fields, they are not significantly attenuated or distorted by tissue conductivity, enabling reliable control of embedded magnetic components even in deep luminal environments. This property makes magnetic actuation particularly suitable for in vivo biomedical applications. Furthermore, magnetic actuation is widely recognised as biologically safe. Previous studies have demonstrated the safety of magnetically controlled catheters and guidewires even under magnetic field strengths as high as 7 T1.\nOperation inside MR/CT environments:\nOur system is currently compatible with imaging modalities such as CT and ultrasound and can be navigated under X-ray fluoroscopic guidance in vivo. Magnetic Resonance Imaging (MRI), however, generates strong magnetic field gradients. We therefore do not recommend using the Helixoft system under MRI guidance at this stage. Combining the system with MRI would require a redesign of the control strategy. One possible approach is to employ the MRI scanner itself for both actuation and imaging of the Helixoft system, as initiated by Sitti’s group (Science Advances, 2023, 9, eadg6438). However, potential safety risks, particularly thermal injury, must be carefully addressed.\nThe discussion of the operation inside the MRI has been updated in the Discussion section (Page 21):\n“In addition, although these catheters are compatible with imaging modalities such as CT and ultrasound, their combination with Magnetic Resonance Imaging (MRI) would necessitate a redesign of the control strategy, such as using the MRI scanner itself for both actuation and imaging59. Potential safety risks, particularly thermal injury, would also need to be carefully addressed.”\nComment 3: In-vivo validation is preliminary and incomplete\nOnly a single 30 kg pig is reported for bronchial drug delivery (Fig. 4), with no quantitative metrics on navigation success rate, dose accuracy, mucosal trauma, or acute inflammatory response.\nResponse to Comment 3:\nWe appreciate the reviewer’s valuable comments regarding the in vivo validation. To strengthen the evidence, we conducted new bronchial interventional procedures in two live pigs (Fig. R6), focusing on quantitative metrics including navigation success rate, dose accuracy, mucosal trauma, and acute inflammatory response.\nFig. R6 Photographs of pigs for two additional bronchial interventional procedures.\nNavigation success rate:\nWe executed ten independent navigation trials toward a target position using our Helixoft-integrated microcatheter. Real-time X-ray guidance and endoscopic imaging are employed to improve accuracy. The microcatheter successfully reaches the predefined target position in all trials (Figs. R7 and R8). The mean duration of navigation from the trachea to the quaternary bronchus is approximately 4 minutes.\nThe description of the navigation success rate has been added to the manuscript (Page 16), as follows:\n“To assess navigation success rate, ten independent navigation trials are performed in total. Under combined real-time X-ray and endoscopic imaging, all trials successfully reach the target position (Supplementary Figs. 10 and 11). The mean navigation time from the trachea to the quaternary bronchus is approximately 4 minutes.”\nFig. R7 Repeated in vivo bronchial navigation experiments in a live pig. a, Stiffness tuning process visualised under simultaneous X-ray and camera imaging. b, Successful repeated navigation to the quaternary bronchus over four independent trials. Scale bar: 10 mm.\nFig. R8 Terminal positions of microcatheter from additional navigation trials. Representative end-point images from six further navigation attempts, each demonstrating precise arrival of the microcatheter at the same quaternary bronchus. Scale bar: 10 mm.\nDose accuracy:\nWe use a Fluigent pressure system to precisely control the amount of drug (Fig. R9). 25 μL of the liquid drug is administered through the microcatheter. After drug injection, an additional volume of 100 μL air (or saline), equivalent to the microcatheter’s internal dead volume (calculated asV ·d /22L , where d is the internal diameter and L is the length)—is injected as a flushing buffer to ensure complete payload delivery. No residual fluid is observed in the catheter, confirming accurate dosing.\nThe relevant content has been added to the manuscript (Page 16) and the Methods section (Page 25), as follows:\nPage 16: “We employ a fluidic control system to precisely regulate the amount of drug (Supplementary Fig. 8). A total of 25 μL of liquid drug is delivered through the microcatheter. The drug distribution is successfully monitored through both the integrated mini-camera and the X-ray radiography in the prone position (Supplementary Fig. 9). After administration, no residual drug remains in the catheter lumen, confirming accurate dosing.”\nPage 25: “A required amount of mock drug was preloaded into the reservoir connected to the microcatheter. The drug solution was then delivered through the microcatheter. Following delivery, an additional 100 μL of air (or saline)—corresponding to the microcatheter’s internal dead volume (V ·d /22L , where d is the internal diameter and L is the length)—was injected as a flushing buffer to ensure complete payload delivery.”\nFig. R9 Drug release control. a, Regulation of drug release using a pressure controller. b, Drug delivery through the microcatheter positioned within the bronchus. c, Physical image of the microcatheter after drug delivery, showing no residual drug.\nMucosal trauma:\nMucosal trauma is evaluated through a combination of on-board camera imaging and histological analysis. Real-time imaging shows no signs of mucosal injury, such as tearing, surface abrasion, or bleeding, during navigation. The histological examination of tissue sections from the primary and tertiary bronchi with H&E staining further confirms an intact mucosal epithelium without apparent mechanical damage.\nFig. R10 Histological analysis of HE-stained tissue sections. a, Schematic illustration of the sampling sites for pathological sections. b, Representative tissue structures observed in the sections. c-d, Pathological analysis results of the HE-stained samples.\nAcute inflammatory response:\nAcute inflammatory response is evaluated through routine clinical blood analyses, including C-reactive protein (CRP), biochemical assays, and complete blood count (CBC). Blood samples are collected before surgery and at 24 h after surgery.\nThe CRP levels in both pigs remain within the normal range (<10 mg/L) before and after the pulmonary intervention (Table R2).\nTable R2. C-reactive protein (CRP) test results\nPreoperative-CRP Postoperative-CRP level\nSample\nlevel(mg/L) (mg/L)\nPig 1 2.8 5.7\nPig 2 4.3 2.6\nReference <10 mg/L: normal\nrange 10–35 mg/L: Mildly elevated (low-grade inflammation)\n35 mg/L: Markedly elevated (significant inflammation)\nBiochemical assay results confirm that liver and kidney functions are unaffected, with no signs of hepatotoxicity or nephrotoxicity (Table R3).\nTable R3. Biochemical assay results\nPig 1- Pig 1- Pig 2- Pig 2- Reference\nParameter\nPreoperative Postoperative Preoperative Postoperative range\nALT (U/L) 42.451 42.885 43.636 43.330 0-95\nAST (U/L) 23.941 24.391 23.882 24.382 16-90\nTP (g/L) 62.159 62.247 63.038 62.720 49-80\nALB (g/L) 32.135 32.043 32.437 32.390 22.3-45.3\nGLB (g/L) 30.024 30.204 30.601 30.330 14-52\nT-Bill\n1.011 0.998 1.352 1.095 0-7.7\n(µmol/L)\nGGT(U/L) 46.445 45.791 48.233 47.865 16-86\nTBA\n17.957 18.261 19.187 19.146 /\n(µmol/L)\nALP(U/L) 116.556 116.075 117.168 116.196 64-576\nBUN\n3.979 3.990 4.189 4.480 2.1-10.7\n(mg/dL)\nCREA\n66.727 64.477 66.537 68.060 44-186\n(µmol/L)\nUa\n6.657 6.664 8.032 7.570 <10.0\n(µmol/L)\nLiver function parameters include ALT (Alanine aminotransferase), AST (Aspartate aminotransferase), GGT (Gamma-glutamyl transferase), TBIL (Total bilirubin), GLOB (Globulin), ALP (Alkaline phosphatase), ALB (Albumin), and TBA (Total bile acid).\nKidney function parameters include BUN (Blood urea nitrogen), UA (Uric acid), and CR (Creatinine).\nIn the CBC analysis (Table R4), most parameters for both pigs fall within the corresponding reference ranges. For pig 2, the white blood cell count (WBC), neutrophil count (Neu#), and neutrophil percentage (Neu%) are slightly below the lower limits of the reference values, while all other indices remain normal.\nTable R4. Complete blood count (CBC) results\nPig 1- Pig 2-\nParameter Pig 1-Preoperative Pig 2-Preoperative Reference range\nPostoperative Postoperative\nWBC (10^9/L) 14.89 15.06 9.81 10.16 11.2-23.1\nNeu# (10^9/L) 7.4 7.49 1.38 1.06 3.1-10.21\nLym#\n5.36 5.46 7.55 8.26 4.8-17.2\n(10^9/L)\nMON #\n1.43 1.43 0.68 0.69 0.4-1.48\n(10^9/L)\nEos#(10^9/L) 0.7 0.68 0.2 0.15 0.00-1.75\nBas#(10^9/L) 0 0 0 0 0.00-0.30\nNeu%（%） 49.7 49.8 14.1 10.5 15.0-63.0\nLym%（%） 36 36.3 76.9 81.3 30.0-82.0\nMON % (%) 9.6 9.4 6.9 6.8 3.0-10.0\nEos%（%） 4.7 4.5 2.1 1.4 0.0-10.0\nBas%（%） 0 0 0 0 0.0-2.0\nRBC\n6.65 6.48 6.05 6.11 5.5-8.4\n(10^12/L)\nHGB (g/L) 122 122 108 108 105-155\nHCT (%) 35.9 34.8 30.4 30.6 33.0-51.8\nMCV (fL) 54 53.8 50.2 50.1 51.0-67.0\nMCH (pg) 18.4 18.9 17.8 17.7 14.5-21.5\nMCHC (g/L) 340 351 355 354 280-340\nRDW-CV (%) 20 20 17.6 17.3 16.0-25.0\nRDW-SD\n39.5 39.5 33.9 33.3 34.0-51.0\n(fL)\nPLT (10^9/L) 295 314 462 465 240-850\nMPV (fL) 9.2 9.2 8.4 8.4 7.8-12.1\nPDW (fL) 16.4 16.3 16 15.9 14.0-17.5\nPCT (%) 0.27 0.289 0.386 0.39 0.210-0.800\nWBC: White Blood Cell Count; Neu#: Neutrophil Count; Lym#: Lymphocyte Count; MON#: Monocyte Count; Eos#: Eosinophil Count; Bas#: Basophil Count; Neu%: Neutrophil Percentage; Lym%: Lymphocyte Percentage; MON%: Monocyte Percentage; Eos%:Eosinophil Percentage; Bas%:Basophil Percentage; RBC: Red Blood Cell Count; HGB: Hemoglobin Concentration; HCT: Hematocrit; MCV: Mean Corpuscular Volume; MCH: Mean Corpuscular; Hemoglobin; MCHC: Mean Corpuscular Hemoglobin Concentration; RDW-CV: Red Cell Distribution Width - Coefficient of Variation; RDW-SD: Red Cell Distribution Width - Standard Deviation; PLT: Platelet Count; MPV: Mean Platelet Volume; PDW: Platelet Distribution Width; PCT: Plateletcrit\nRoutine clinical blood analysis results show no significant postoperative changes, demonstrating the absence of acute inflammation or organ dysfunction. Table R5 summarises the assessment results, including navigation success rate, drug delivery accuracy, mucosal trauma, and acute inflammatory response. Together, these evaluations confirm both the efficacy and biosafety of our system.\nTable R5. Assessment results of pulmonary interventions in live pig models\nParameter Pig 1 Pig 2\nNavigation success rate 100 % (avg. 4 min, 10 trails) 100 % ((avg. 4 min, 10 trails)\nDrug delivery accuracy No visible drug residue in catheter No visible drug residue in catheter\nMucosal trauma by camera imaging None observed (erythema, spasm, None observed (erythema, spasm,\nbleeding) bleeding)\nAcute inflammatory response (CRP, None None\nbiochemical assays, and CBC)\nHistological analysis Intact mucosal epithelium Intact mucosal epithelium\nPost-operative observation:\nBoth pigs were observed for 7 days after surgery and showed no abnormal feeding, sleeping, or physical activities (Fig. R11). Body weight was maintained throughout the observation period. These findings confirm the biocompatibility and safety profile of our system.\nFig. R11 Postoperative monitoring of two pigs over a 7-day period. The bottom-right chart in each panel displays the corresponding body weight.\nThe assessment results are presented in the main text (page 16), Supplementary Note 3, Supplementary Tables 3–6, and Supplementary Figs. 12–13.\nPage 16 of manuscript: “Post-operative assessment confirms the safety profile of the intervention, as indicated by maintenance of stable body weight and normal activity in the porcine models over a 7-day post-operative observation period (Supplementary Fig. 12). Furthermore, routine clinical blood analyses and histological examinations reveal no signs of acute inflammation or mucosal damage. (Supplementary Fig. 13, Supplementary Note 3, and Supplementary Tables 3–6).”\nPages 11 and 12 of Supplementary Information:\n“Supplementary Note 3 | Systemic safety assessment of the Helixoft-integrated microcatheter\nTo verify the safety of pulmonary interventional surgery using our system, we perform routine clinical blood analyses and histological examinations. The acute inflammatory response is evaluated through routine clinical blood analyses, including C-reactive protein (CRP), biochemical assays, and complete blood count (CBC). Blood samples are collected before surgery and at 24 h after surgery. The CRP levels in both pigs remained within the normal range (<10 mg/L) before and after the pulmonary intervention (Supplementary Table 3). Biochemical assay results confirm that liver and kidney functions are unaffected, with no signs of hepatotoxicity or nephrotoxicity (Supplementary Table 4). In the CBC analysis (Supplementary Table 5), most parameters for both pigs fall within the corresponding reference ranges. For pig 2, the white blood cell count (WBC), neutrophil count (Neu#), and neutrophil percentage (Neu%) were slightly below the lower limits of the reference values, while all other indices remained normal. Routine clinical blood analysis results show no significant postoperative changes, demonstrating the absence of acute inflammation or organ dysfunction. Mucosal trauma is evaluated through a combination of on-board camera imaging and histological analysis. Real-time imaging shows no signs of mucosal injury, such as tearing, surface abrasion, or bleeding, during navigation. The histological examination of tissue sections from the primary and tertiary bronchi with H&E staining further confirms an intact mucosal epithelium without apparent mechanical damage (Supplementary Fig. 13). Supplementary Table 6 summarises the assessment results, including navigation success rate, dose accuracy, mucosal trauma, and acute inflammatory response. Together, these evaluations confirm the efficacy and safety of our system.”\nComment 4: Modelling assumptions may not capture critical nonlinearities\nThe constant-curvature beam model (Eq. 1) presumes a uniform EI and ignores the locally stiff/free transition that occurs as the helix overlaps the soft tube. This simplification explains the ≤10 % MAPE reported, but those tests were run in air with negligible distributed contact forces. In tortuous lumina, distributed wall contact introduces piecewise-loading conditions; without contact mechanics, the model’s predictive power for in-situ steering remains unproven.\nResponse to comment 4:\nWe thank the reviewer for pointing out the limitations of the constant-curvature beam model. On-board force sensing for microcatheter and guidewire, particularly at the distal end, is a long-standing challenge in the field of minimally invasive surgery. The steering of our Helixoft system relies on the integration of magnetic torque and real-time imaging guidance. In our initial model, we did not explicitly consider distributed wall contact, as it mainly generates tangential friction along the catheter wall, as demonstrated in our friction hysteresis analysis (Response to Comment 2). Under X-ray fluoroscopy or other imaging guidance modalities, tortuous lumens requiring in-situ steering can be clearly visualised. The rapid magnetic responsiveness of our Helixoft system enables real-time control of catheter curvature, minimising excessive tip–wall contact and allowing precise navigation.\nWe acknowledge, however, that the reviewer's comment regarding the influence of distributed contact forces is both insightful and constructive. We agree that a suitable modelling approach would not only be theoretically beneficial for the development of more autonomous systems in our future work, but also provide valuable insights for research on magnetically controlled catheters. Consequently, we have expanded our model to incorporate the effects of local stiff–soft transitions and distributed contact forces, as follows:\nPiecewise stiffness model:\nTo analyse the effect of local stiff–soft transitions, a piecewise stiffness model is employed. The microcatheter is divided into rigid and soft segments. The bending stiffness of each segment is calculated as follows:\nFor the rigid segment:\nEI E I E I (5)\n1 1 helix helix ecoflex ecoflex\n \nIn which, I  （d 4-d 4）, I  （d 4-d4）, d is the outer diameter of\nhelix 64 o m1 ecoflex 64 m2 i o\nthe helix, d is the inner diameter of the helix, d is the outer diameter of the soft\nm1 m2\nmicrotube, and d is the inner diameter of the soft section.\ni\nThe bending stiffness of the soft segment can be expressed as:\nE I E I (6)\n2 2 ecoflex ecoflex\nBased on these relationships, the bending angles of each segment can be expressed as follows. For the rigid section, the bending angle is:\nLMB sin（r-）\n  2 1 (7)\nLMB sin（r-）\n  2 2 (8)\n (9)\n1 2\nThe tip pose can be determined as follows.\n EI E I\nx 1 1 sin 2 2 (sin()sin)(L L )cos\n  MB sin（r-） 1 MB sin（r-） 1 2 1 G M\n2 1 2 2\n\nEI E I\n y 1 1 (1cos) 2 2 (coscos())(L L )sin\n MB sin（r-） 1 MB sin（r-） 1 1 2 G M\n\n2 1 2 2\n(10)\nInfluence of distributed contact forces on bending：\nTo address in-situ behaviour, we further consider the influence of distributed contact along the lumen. The net moment acting on the distal end is expressed as:\nL\nM M  q (s)l ds (11)\nnet c lever\nc lever\ncorresponding moment arm.\ndy\nq (s)k (s)y(s)c (12)\nc c dt\nIn which, cis the fluid viscous damping coefficient, k (s) is the spatial distribution\nc\nof tissue stiffness, y(s) is the lateral displacement of the guidewire at contact position.\nThe effect of the distributed contact force on the bending angle can be expressed as：\nL 1 S\n  ( q (s)(s)d)ds (13)\nThe discussions of the piecewise stiffness model and distributed contact force have been added to the Supplementary Information (Pages 9-11), as follows:\n“(3) piecewise stiffness model:\nTo analyse the effect of local stiff–soft transitions, a piecewise stiffness model is employed. The microcatheter is divided into rigid and soft segments. The bending stiffness of each segment is calculated as follows:\nFor the rigid segment:\nEI E I E I (15)\n1 helix helix ecoflex ecoflex\n \nIn which, I  （d 4-d 4） and I  （d 4-d4）.\nhelix 64 o m1 ecoflex 64 m2 i\nThe bending stiffness of the soft segment can be expressed as:\nEI E I (16)\n2 ecoflex ecoflex\nBased on these, the bending angles of each segment are:\nFor the rigid section, the bending angle is:\nLMB sin（r-）\n  2 1 (17)\n1 1\nFor the soft section, the bending angle is:\nLMB sin（r-）\n  2 2 (18)\n2 2\nThe total bending angle can be expressed as:\n (19)\n1 2\nThe tip pose can be obtained using planar arc geometry.\n EI E I\nx 1 1 sin 2 2 (sin()sin)(L L )cos\n  MB sin（r-） 1 MB sin（r-） 1 2 1 G M\n2 1 2 2\n\nEI E I\n y 1 1 (1cos) 2 2 (coscos())(L L )sin\n MB sin（r-） 1 MB sin（r-） 1 1 2 G M\n\n2 1 2 2\n(20)\n(4) Influence of distributed contact forces on bending:\nTo address in-situ behaviour, we further consider the influence of distributed contact along the lumen21. The net moment acting on the distal end is expressed as:\nL\nM M  q (s)l ds (21)\nnet c lever\nc\ndy\nq (s)k (s)y(s)c (22)\nc c dt\nThe effect of the distributed contact force on the bending angle can be expressed as：\nL 1 S\n  ( q (s)(s)d)ds (23)”\nComment 5: Minor comments\nTerminology – The term “Helixoft” is catchy but potentially confusing; please clarify in the Abstract that it is not a brand name but a lab prototype.\nResponse to comment 5: Thanks for the reviewer’s valuable suggestion on the term “Helixoft”. We have added a clarification in the Abstract, as follows:\n“(Helixoft, a lab prototype)”\nComment 6: Transition speed – The helix translates at ≈5 mm s⁻¹ (Fig. 2d); with a 10 mm soft segment, a full soft→stiff transition therefore requires ~2 s plus magnetic-arm repositioning time. Explicitly compare this with the <300 ms transition reported for fiber-jamming catheters.\nResponse to comment 6:\nWe thank the reviewer for the thoughtful comment regarding transition speed, as also suggested in Comment 12. Although the fibre-jamming approach achieves a sub-300 ms transition time, its device diameter is limited to ~2.3 mm due to the integration of ~1000 fibres9. In addition, this design lacks a working channel for biopsy or payload delivery. Consequently, fibre-jamming catheters have been primarily developed for specific cardiovascular applications.\nIn narrower lumens, including bronchioles, cerebral vasculature and fallopian tube, device dimensions and discrete stiffness states may impose certain limitations. In these scenarios, conventional surgical approaches require tens of minutes22,23. Moreover, such approaches lack stiffness modulation capability and cannot access the narrower lumens. These challenges highlight the urgent need for a miniaturised, safe, and reliable variable-stiffness catheter, rather than striving for millisecond-scale transition speeds.\nTo address this need, we propose Helixoft—a helix-shaped magnetic soft microrobot that enables in situ and continuous stiffness tuning within microcatheters as small as 300 μm in diameter, without reliance on potentially harmful stimuli.\nIn addition, the Helixoft system allows rapid tuning by increasing the magnetic field frequency to 20 Hz, which shortens the transition time and enables 10-fold tuning in stiffness within 0.3 s. However, this rapid tuning mode is generally avoided in practice due to stability concerns.\nThe discussion of the transition speed has been added in the Discussion section (Page 20), as follows:\n“In addition, the Helixoft system allows rapid tuning by increasing the magnetic field frequency to 20 Hz, which shortens the transition time and enables 10-fold tuning in stiffness within 0.3 s. However, this rapid tuning mode is generally avoided in practice due to stability concerns.”\nComment 7: Lubricious coating durability – Long-term adhesion and cytotoxicity of the PDMAA coating are not characterised. Provide peel/abrasion data over >100 bend cycles in saline at 37 °C.\nResponse to comment 7: We appreciate the reviewer’s insightful comment concerning the long-term durability and biocompatibility of the PDMAA coating.\nWe conducted a new experiment to evaluate the cytotoxicity of the PDMAA coating. The hydrogel-coated microcatheter is co-cultured with cells (BEAS-2B) for 24 hours (Fig. R12)16. Fig. R12 demonstrates high cell viability, with cells spreading uniformly around the microcatheter, indicating that the hydrogel coating does not pose cytotoxic effects.\nFig.R12 Cytotoxicity assessment of the microcatheter with hydrogel coating via cell co-culture. a, Photograph of the microcatheter co-cultured with cells. b, Fluorescence image showing live (green) and dead (red) cells after 24 hours of incubation. Scale bar: 100 μm.\nTo further assess the long-term adhesion of the PDMAA coating, a thin fluorescein-loaded hydrogel layer is first applied onto bulk Ecoflex to confirm the feasibility of hydrogel adhesion (Fig. R13). The same hydrogel is then successfully coated onto the microcatheter surface. Long-term adhesion and peel/abrasion resistance are evaluated by subjecting the hydrogel-coated microcatheter to 100 bending cycles in 0.9% saline at 37 °C. The bending primarily occurs in the soft microtube section. Following the bending test, the hydrogel coating is immersed in a fluorescent dye solution. As shown in Fig. R13, the fluorescence signals after bending verify that the hydrogel remained securely attached throughout bending, with no signs of significant peeling or abrasion.\nFig. R13 Stability of hydrogel coating during bending. a, Fluorescent image showing the hydrogel skin coated on Ecoflex, visualised using fluorescein. b, Long-term adhesion test of the hydrogel coating after 100 bending cycles.\nThe results of the long-term adhesion and cytotoxicity of the hydrogel coating have been added in Supplementary Figs. 16 and 17. The corresponding content has been updated in the Methods section (Page 23) and Supplementary Note 4 (Page 12) as follows:\nPage 23 of manuscript: “Cytotoxicity and long-term adhesion of the hydrogel coating were evaluated by cell culture and by bending-cycle tests in saline, respectively (Supplementary Note 4 and Supplementary Figs. 16–17).”\nPage 12 of Supplementary Information:\n“Supplementary Note 4 | Long-term adhesion and cytotoxicity of the hydrogel coating\nTo assess cytotoxicity, the hydrogel-coated microcatheter is co-cultured with cells (BEAS-2B) for 24 hours8. Supplementary Fig.16 demonstrates high cell viability, with cells spreading uniformly around the microcatheter, indicating that the hydrogel coating does not pose cytotoxic effects. To further assess the long-term adhesion of the PDMAA coating, a thin fluorescein-loaded hydrogel layer is first applied onto bulk Ecoflex to confirm the feasibility of hydrogel adhesion (Supplementary Fig.17). The same hydrogel is then successfully coated onto the microcatheter surface. Long-term adhesion and peel/abrasion resistance are evaluated by subjecting the hydrogel-coated microcatheter to 100 bending cycles in 0.9% saline at 37 °C. The bending primarily occurs in the soft microtube section. Following the bending test, the hydrogel coating is immersed in a fluorescent dye solution. As shown in Supplementary Fig.17, the fluorescence signals after bending verify that the hydrogel remained securely attached throughout bending, with no signs of significant peeling or abrasion.”\nComment 8: Magnetic safety – 150 mT static fields within 3–5 cm of the thoracic cavity warrant a discussion of pacemaker/ICD compatibility.\nResponse to Comment 8:\nWe thank the reviewer for raising this important point. While our magnetically controlled catheters show great potential in biopsy and payload delivery, their use in patients who have magnetic implants, such as cardiac pacemakers or implantable cardioverter-defibrillators (ICDs), require careful planning due to safety concerns related to the interaction of these devices with magnetic fields. A clarification has been added in the Discussion section (Pages 20 and 21), as follows:\n“While magnetically controlled catheters show great potential in biopsy and payload delivery, their use in patients who have magnetic implants, such as cardiac pacemakers or implantable cardioverter-defibrillators (ICDs), require careful planning due to safety concerns related to the interaction of these devices with magnetic fields59.”\nComment 9: Serilisation and re-use – The assembly includes stainless-steel helix + silicone tube glued to a PE catheter. How does ethylene-oxide or plasma sterilisation affect the glue joint and magnetic remanence?\nResponse to comment 9:\nWe thank the reviewer for raising this important question regarding sterilisation. We conducted new experiments to evaluate the effect of ethylene oxide (EtO) sterilisation on the glue joint integrity and the magnetic remanence of the Helixoft system. A schematic of the test procedure is shown in Fig. R14. After sterilisation, the structure remains intact, and the magnetic remanence of the components is stable. We further confirm the structural integrity by performing a repeated bending test (n = 100 cycles) and payload delivery.\nWe have added the test procedure in Supplementary Fig. 15 of the Supplementary Information, and details of the sterilisation of the microcatheter have been added in the Methods section (Page 23), as follows:\n“The microcatheter was sterilised using ethylene oxide (EtO). This sterilisation process did not affect its magnetisation or structural integrity, as verified by surface magnetic flux intensity measurements, repeated bending tests, and payload delivery performance (Supplementary Fig. 15).”\nFig. R 14 Sterilisation of the Helixoft-integrated microcatheter: a 1-mm-diameter device with biopsy needle and magnetic tubular tip, as well as a 300-μm microcatheter with a magnetic tubular tip. a, Schematic of the test procedure for evaluating the effects of sterilisation on structural integrity and magnetic remanence. b, Photograph of the sterilised and packaged devices. c, Results of structural integrity testing and surface magnetic flux density measurements.\nComment 10: Nomenclature of stiffness states – “Low, medium, high” states correspond to specific helix rotations (0, 3, 6 turns), but the absolute EI values vary with catheter diameter. Report both the helix pitch coverage and the resulting EI to avoid ambiguity.\nResponse to comment 10:\nWe thank the reviewer for this valuable suggestion. In the revised manuscript, the stiffness states are now defined by helix pitch coverage length (L₁=turns×pitch) together with the corresponding bending stiffness (EI), thereby avoiding ambiguity related to catheter diameter. Specifically, for the helix with a pitch of 1 mm and diameter of 1 mm, the three states are: Low stiffness: (pitch coverage L₁= 0 mm, EI=2.06×10⁻⁸ N·m²), Medium stiffness:( pitch coverage L₁=3 mm, EI=3.60×10⁻⁸ N·m²), and High stiffness: (pitch coverage L₁ =6 mm, EI=1.72×10⁻⁷ N·m²). For the helix with a pitch of 0.2 mm and diameter of 1 mm, the three states are: Low stiffness: (pitch coverage L₁=0 mm, EI=1.72×10⁻9 N·m²), Medium stiffness:( pitch coverage L₁=1.9 mm, EI=9.72×10⁻9 N·m²), and High stiffness: (pitch coverage L₁ = 3.8 mm, EI= 7.43×10⁻8 N·m²).\nThe details of the stiffness states have been added in the Methods section (Page 23), as follows:\n“The stiffness states were defined by helix pitch coverage length (L₁=turns×pitch) together with the corresponding bending stiffness (EI), thereby avoiding ambiguity related to catheter diameter. Specifically, for the helix with a pitch of 1 mm and diameter of 2 mm, the three states were: Low stiffness: (pitch coverage L₁=0 mm, EI=2.06×10⁻⁸ N·m²), Medium stiffness:( pitch coverage L₁=3 mm, EI=3.60×10⁻⁸ N·m²), and High stiffness: (pitch coverage L₁=6 mm, EI=1.72×10⁻⁷ N·m²). For the helix with a pitch of 0.2 mm and diameter of 1 mm, the three states were: Low stiffness: (pitch coverage L₁=0 mm, EI=1.72×10⁻9 N·m²), Medium stiffness:( pitch coverage L₁=1.9 mm, EI=9.72×10⁻9 N·m²), and High stiffness: (pitch coverage L₁=3.8 mm, EI=7.43×10⁻8 N·m²).”\nComment 11: To reach the novelty and depth expected for Nature Communications, the authors should:\n1. Demonstrate a capability that prior devices cannot match—e.g., real-time decoupling under simultaneous R- and S-field actuation inside tissue-mimicking phantoms with distributed constraints, or a statistically significant performance gain (>2×) in task-completion time over a leading benchmark catheter.\nResponse to comment 11:\nWe thank the reviewer for the valuable suggestion concerning the real-time decoupling of R- and S-field actuation as well as the transition time. Details on real-time decoupling are included in our response to Comment 2, and the discussion of transition time is presented in our response to Comment 6.\nComment 12: 2. Provide a rigorous, quantitative benchmarking table that covers stiffness range, transition time, bending angle per Tesla, force throughput, and device OD across at least five state-of-the-art systems.\nResponse to comment 12:\nWe thank the reviewer for the valuable suggestion regarding the quantitative benchmarking table. As suggested, we have added a quantitative comparison table that covers stiffness range, transition time, bending angle per Tesla, force throughput, and device outer diameter (OD) across representative state-of-the-art systems (Table R1). We also summarise additional parameters, such as safety and remote controllability, and in vivo verification and quantitative safety assessment\nThe Helixoft system achieves an outer diameter of 300 µm, enabling application in narrow lumens (e.g., bronchioles, cerebral vasculature and fallopian tube). It employs a single-field decoupling strategy for stiffness tuning and steering via magnetic actuation, thereby ensuring safety and remote controllability. We validated the feasibility and safety of the system in pulmonary interventional procedures on three live porcine models, with quantitative safety assessments including navigation success rate, histological evaluation, and acute inflammatory response.\nComment 13: 3. Expand in-vivo validation to at least n = 3–5 large-animal procedures with histological assessment and device-specific adverse-event reporting.\nResponse to comment 13:\nWe appreciate the reviewer’s valuable comment regarding the number of large-animal experiments and the importance of thorough safety evaluation. As shown in Figs. R6-R9 and the answer to Comment 3, we expanded the in vivo validation by performing two additional large-animal trials, resulting in a total of three trials (n = 3). We assessed the system’s efficacy through navigation success rates and drug delivery precision. We further evaluated safety, including mucosal trauma, acute inflammatory responses, and post-operative observation. Collectively, these results strengthen the evidence for both the efficacy and safety of the Helixoft system in vivo.\nReviewer #3\nComment 1: This paper describes a microrobotic system that uses a helix-shaped magnetic soft microrobot positioned around a commercially catheter. Soft microrobot enabling continuous stiffness tuning and precise steering. The approach is interesting however there are some inconsistences. That I will describe in detail as follow:\n1. Figure 1 is a diagram of a person inside of a 3-coil arrangement. The validation of the system was performed in an ex-vivo oviduct biopsy and in-vivo drug delivery to the bronchi of a live pig. In this way the scheme in figure 1 is not consistent with the experiments.\nResponse to comment 1:\nWe thank the reviewer for the kind suggestion regarding the caption of Figure 1. Figure 1 is indeed a conceptual illustration of the applicable scenarios of our microrobotic system, rather than a depiction of the experimental process. As suggested by the reviewer, we have revised the caption of Figure 1 accordingly as:\n“Conceptual illustration of the Helixoft microrobotic system for stiffness tuning and active steering of microcatheters during interventional surgeries.”\nComment 2: 2. To generate the magnetic actuation in this paper they use a 3 coils system and a permanent magnet mounted on a robotic arm. Usually the rotating magnetic field is generated by a Helmholtz coil array consisting of three pairs of coils or something similar. This can of system con perfectly control the direction of magnetic field, why they use a permanent magnet mounted on a robotic arm. A detailed description of the design and operation of the magnetic actuation systems should be included in the Materials and Methods section, as well as photographs in the Supporting Information.\nResponse to comment 2:\nWe thank the reviewer for this suggestion. As suggested by the reviewer, we have added a detailed description of the control module and the setup of the Methods section (Page 23), as shown below:\n“The magnetic control module employed either a coil system or a permanent magnet mounted on a robotic arm. The coil-based module consisted of three pairs of custom-built Helmholtz coils made of copper wire and connected to a direct current (DC) power supply. The three pairs of coils generated a three-axis magnetic field with 99% field homogeneity, producing field strengths up to 30 mT within a 50-mm spherical region over a frequency range of 0-20 Hz. The magnet-based module used a permanent block magnet (45 mm×45 mm×21 mm) mounted on a robotic arm, providing a maximum field strength of 200 mT.”\nComment 3: 3. Importantly, the paper lacks crucial information regarding the experiments. For example, there's no detailed description of the magnetic controller and its functionality, nor is the material composition and fabrication method of the rigid helical microrobot. Seems that they use stainless steel with magnetic properties… The authors need to include this information in the materials and methods section.\nResponse to comment 3:\nWe appreciate the reviewer’s insightful comment on the magnetic controller and the fabrication material of the rigid helix. The relevant methodological details have been added to the Methods section (Pages 22 and 24):\nPages 22: “The rigid magnetic helix was fabricated by tightly winding a stainless-steel wire around a cylindrical metal rod, where the helix diameter was determined by the rod’s diameter. The pitch and number of turns were adjusted to tune the stiffness and length of the structure. The rigid magnetic helix, with higher stiffness, was coaxial with the soft microtube, which had lower stiffness. Three rigid magnetic microhelices with varying dimensions—0.5 mm (0.08-mm wire diameter, 0.3-mm pitch), 1.1 mm (0.2-mm wire diameter, 0.2-mm pitch), and 2 mm (0.1-mm wire diameter, 1-mm pitch)—were matched individually with corresponding soft microtubes.”\nPages 24: “The robotic arm with six revolute joints was used to manoeuvre the cube magnet. The magnetic control module generated the R- and S-fields for stiffness tuning and active steering. In the coil-based module, field strength and frequency were adjusted by varying the coil current, with parameters controlled through a graphical user interface (GUI). In the magnet-based module, the R- and S-fields were produced by rotating or reorienting the magnet with robotic arm control, and the R-field speed was set by the motor rotation frequency (0–5 Hz). The coil-based module was mainly used for benchtop experiments and 3D phantom models, whereas the magnet-based module was applied in vivo to provide a larger working space. To accommodate different penetration depths in vivo, magnets of varying sizes and field strengths were selected as required.”\nComment 4: 4. The paper fails to describe the Helixoft microrobot's morphology characterization. Crucially, it omits information regarding the composition of its magnetic components (percentage and type) and the magnetization curve of its ferromagnetic material.\nResponse to Comment 4:\nWe thank the reviewer for pointing out the need to clarify the morphology and magnetic characterisation of the Helixoft microrobot. In order to demonstrate the magnetic characterisation of magnetic components, we implemented vibrating sample magnetometry (VSM) measurement of magnetic components (Fig. R 14).\nThe composition and fabrication of the tubular distal end and these magnetisation curves have been added in the Methods section on Page 21, as follows:\n“The tubular distal end was modular and material-selective, tailored to specific functional requirements. For tissue biopsy, a stainless-steel needle structure (outer diameter: 0.5 mm) was employed and partially filled with an NdFeB–Ecoflex composite (10 wt% NdFeB particles) to ensure magnetisation stability. The delivery channel for payloads was defined by a fibre inserted during the composite filling process, resulting in a channel approximately 100 μm in diameter. For applications without biopsy requirements, the distal head was fully enclosed within the microcatheter. All magnetic components were magnetised using a high-field magnetizer, with five magnetization cycles to ensure long-term stability of magnetisation60,61. The magnetic components (stainless steel helix, biopsy needle with NdFeB–Ecoflex composite, ring-shaped NdFeB magnet, and pure biopsy needle) were characterised using vibrating sample magnetometry (VSM) (Supplementary Fig. 14).”\nFig. R15 Magnetisation curves of the magnetic components.\nReference\n1. Tiryaki, M. E., Elmacioǧilu, Y. G. & Sitti, M. Magnetic guidewire steering at ultrahigh magnetic fields. Sci. Adv. 9, eadg643 (2023).\n2. Li, Z. et al. Design and Hierarchical Control of a Homocentric Variable-Stiffness Magnetic Catheter for Multiarm Robotic Ultrasound-Assisted Coronary Intervention. IEEE Trans. Robot. 40, 2306–2326 (2024).\n3. Richter, M., Venkiteswaran, V. K. & Misra, S. Concentric Tube-Inspired Magnetic Reconfiguration of Variable Stiffness Catheters for Needle Guidance. IEEE Robot. Autom. Lett. 8, 6555–6562 (2023).\n4. Wang, C. et al. Heterogeneous multiple soft millirobots in three-dimensional lumens. Sci. Adv. 10, eadq1951 (2024).\n5. Abbott, J. J., Ergeneman, O., Kummer, M. P., Hirt, A. M. & Nelson, B. J. Modeling magnetic torque and force for controlled manipulation of soft-magnetic bodies. IEEE Trans. Robot. 23, 1247–1252 (2007).\n6. Beleggia, M., Vokoun, D. & De Graef, M. Demagnetization factors for cylindrical shells and related shapes. J. Magn. Magn. Mater. 321, 1306–1315 (2009).\n7. Fu, S. et al. A Magnetically Controlled Guidewire Robot System with Steering and Propulsion Capabilities for Vascular Interventional Surgery. Adv. Intell. Syst. 5, 2300267 (2023).\n8. Li, J. & Wang, L. Modeling magnetic soft continuum robot in nonuniform magnetic fields via energy minimization. Int. J. Mech. Sci. 282, 109688 (2024).\n9. Sun, Y. et al. Instant variable stiffness in cardiovascular catheters based on fiber jamming. Sci. Adv. 11, eadn1207 (2025).\n10. Lussi, J. et al. A Submillimeter Continuous Variable Stiffness Catheter for Compliance Control. Adv. Sci. 8, 2101290 (2021).\n11. Yang, P., Mao, L., Tian, C., Meng, X. & Xie, H. A Cooperative and Multifunctional Magnetic Continuum Robot for Noninteractive Access, Dexterous Navigation, and Versatile Manipulation. Adv. Funct. Mater. 35, 2412543 (2024).\n12. Chautems, C. et al. Magnetic Continuum Device with Variable Stiffness for Minimally Invasive Surgery. Adv. Intell. Syst. 2, 1900086 (2020).\n13. Piskarev, Y. et al. Fast-Response Variable-Stiffness Magnetic Catheters for Minimally Invasive Surgery. Adv. Sci. 11, 2305537 (2024).\n14. Piskarev, Y. et al. A Variable Stiffness Magnetic Catheter Made of a Conductive Phase-Change Polymer for Minimally Invasive Surgery. Adv. Funct. Mater. 32, 202107662 (2022).\n15. Kim, J., Choi, W. Y., Kang, S., Kim, C. & Cho, K. J. Continuously Variable Stiffness Mechanism Using Nonuniform Patterns on Coaxial Tubes for Continuum Microsurgical Robot. IEEE Trans. Robot. 35, 1475–1487 (2019).\n16. Mao, L. et al. Magnetic steering continuum robot for transluminal procedures with programmable shape and functionalities. Nat. Commun. 15, 3759 (2024).\n17. Mattmann, M. et al. Thermoset Shape Memory Polymer Variable Stiffness 4D Robotic Catheters. Adv. Sci. 9, 202103277 (2022).\n18. Liu, H. et al. A concentric tube magnetic continuum robot with multiple stiffness levels and high flexibility for potential endovascular intervention. J. Magn. Magn. Mater. 597, 172023 (2024).\n19. Schenck, J. F. MR safety at high magnetic fields. Magn. Reson. Imaging Clin. N. Am. 6, 715–730 (1998).\n20. Nath, M. N., Ulvog, A. K., Weigand, S. & Sen, S. Understanding the Role of Magnetic and Magneto-Quasistatic Fields in Human Body Communication. IEEE Trans. Biomed. Eng. 69, 3635–3644 (2022).\n21. de Silva, C. W. Mechanics of Materials. Mechanics of Materials (2013). doi:10.1201/b15431.\n22. Lee, E. J. et al. Procedure Time of Endovascular Thrombectomy as Performance Measure of Acute Stroke Treatment. Neurointervention 20, 71–81 (2025).\n23. VanderLaan, P. A., Wang, H. H., Majid, A. & Folch, E. Endobronchial ultrasound-guided transbronchial needle aspiration (EBUS-TBNA): An overview and update for the cytopathologist. Cancer Cytopathology vol. 122 561–576 (2014).\nReviewer #2\nComment 1: I am satisfied with the substantial improvements in the revised manuscript. The added model depth, inclusion of animal studies, and overall rigor represent a marked advance over the initial submission.\nResponse to comment 1:\nWe sincerely thank the reviewer for the encouraging comments and for recognizing the improvements made in the revised manuscript. We truly appreciate the reviewer’s positive assessment of the added model depth, inclusion of animal studies, and the overall enhancement in rigor.\nComment 2: However, the authors have not addressed my primary comment:\n“Major comments\nLimited conceptual novelty relative to prior art on magnetic, variable-stiffness catheters\nThe manuscript presents a helix-sliding sleeve that shortens/elongates a soft segment to achieve a 40-fold stiffness change, coupled with radial-magnet torque steering. Yet the key ingredients (i) a magnetically actuated rigid helix, (ii) sliding-sleeve stiffness modulation, and (iii) decoupled torque steering each appear in the literature individually or in combination. Representative examples include LMPA-based or SMP-based magnetic catheters with multi-segment stiffness switching (Chautems et al., 2020; Piskarev et al., 2023) as well as continuum robots that integrate ferromagnetic steering with mechanically programmable stiffness (Kim et al., 2019; Mao et al., 2024).\nThe manuscript, therefore, risks being seen as an incremental combination rather than a substantive advance.”\nResponse to comment 2:\nWe sincerely thank the reviewer for highlighting this important point regarding the conceptual novelty of our work. To address this concern, we provide a detailed comparison between our Helixoft-integrated microcatheter system and previously reported magnetic catheters with stiffness-modulation capabilities in Table R1.\nWhile these earlier studies each demonstrated distinct advantages, they also face significant limitations, particularly in terms of device miniaturization, operational safety, and remote controllability. For instance, LMPA-based or SMP-based magnetic catheters with multi-segment stiffness switching (Chautems et al., 2020;1 Piskarev et al., 20232 ) can achieve a broad stiffness modulation range (> 66-fold). However, their dependence on intense thermal or electrical stimuli raises safety concerns in delicate luminal environments, including potential material toxicity, thermal injury, and electrical damage. Moreover, the reported minimum outer diameters of these devices remain around 2.3 mm, as the intrinsic constraints of their actuation mechanisms hinder further miniaturization and integration at smaller scales. Similarly, concentric-tube-based stiffness-modulation catheters with continuous switching capability (Kim et al., 2019)3 face persistent challenges in downsizing due to the complex assembly of proximally controlled multilayer continuum tubes. The follow-the-leader catheter employing LMPA-based stiffness modulation (Mao et al., 2024)4 achieves periodic, tip-based elongation under programmable magnetic fields and enables autonomous shaping, yet the coaxial configuration of dual tubes results in device diameters of 3–4 mm, restricting their use in narrow lumens.\nIn contrast, the present Helixoft system is specifically designed for navigation within narrow and sensitive luminal environments (e.g., bronchioles, cerebral vasculature, and fallopian tubes) by miniaturizing the variable-stiffness catheter to 300 μm in diameter. Achieving tunable mechanical performance at the microscale remains a major challenge as miniaturization inherently conflicts with the complex structural requirements for stiffness tuning. Our microrobotic system addresses this challenge based on a previously underutilized mechanism that harnesses microscale helical motion at the component level to dynamically modulate the otherwise static mechanical properties of the overall structure under a magnetic field. Through a decoupling magnetic control strategy, independent remote tuning of stiffness and active steering is achieved by combining helical motion with torque-driven bending of individual microrobotic components. This mechanism eliminates the need for potentially harmful stimuli such as thermal, optical, or electrical inputs. Furthermore, repeated in vivo validations across different animal models, together with systematic safety evaluations, confirm both the feasibility and biocompatibility of the Helixoft system. These core advantages are summarized in Table R1.\nTable R1. Comparison of the Helixoft-integrated microcatheter with existing variable-stiffness and steerable catheters.\nDevice type Device Safe and Continuous In vivo Transition time stiffness Bending Force\nouter remote stiffness validation Softening Stiffening r ange angle per throughput\ndiameter control & safety Tesla\n≤1 mm\nMagnetically controlled 300-1000 Yes Continuous Yes(3 cases 0.3 s 0.3 s 40-fold 787 1.5 N\nhelical motion (This work) μm (0.12 kPa/dewith safety (118°/150\ng) assessment) mT)\nThermal-responsive low-1 mm No Stepwise No 16 s 30 s 400-fold N/A 8 N\nmelting-point alloy5\n1-3 mm in device diameter\nVacuum-responsive fibre 2.3 mm No Stepwise No 0.1 s 0.3 s 6.5-fold 1337(107°N/A\njamming6 /80 mT)\nThermal-responsive low-2.8/5.7 No Stepwise No N/A N/A N/A N/A N/A\nmelting-point alloy7 mm\nThermal-responsive low-2.33 mm No Stepwise No 15 s 80 s N/A N/A N/A\nmelting-point alloy1\nThermal-responsive shape 2.3 mm No Stepwise No 13 s 4.4 s 66-fold 1587(127°N/A\nmemory polymer2 /80 mT)\nThermal-responsive 2 mm No Stepwise No 175 s 107 s 21-fold 637(51°/8 N/A\nconductive shape memory 0 mT)\npolymer8\nConcentric tube assembly3 1.8 mm No Continuous No N/A N/A 7.2-fold N/A N/A\n3-5 mm in device diameter\nThermal-responsive low-~3-4 mm No Stepwise Yes(1 case 10 s 15 s N/A N/A N/A\nmelting-point alloy4 without\nsafety\nassessment)\nConcentric tube assembly9 3.5 mm No Stepwise No N/A N/A 14-fold N/A ~2.5 N\nThermoset shape memory 3 mm No quasi- No 9 s 60 s 93-fold 2250(90°/ N/A\npolymer10 continuous 80 mT)\n>5 mm in device diameter\nConcentric tube assembly11 7 mm No Stepwise No N/A N/A N/A 6750(135°N/A\n/20 mT\nNotes: Compared with existing variable-stiffness and steerable catheters, our Helixoft-integrated system is downsized to 300 μm and enables continuous stiffness tuning, making it suitable for narrow and sensitive lumens such as the bronchioles, cerebral vasculature, and fallopian tube. Moreover, the system achieves decoupling between stiffness tuning and active steering, both realised under a single magnetic field without introducing any additional potentially harmful stimuli.\nReference\n1. Chautems, C. et al. Magnetic Continuum Device with Variable Stiffness for Minimally Invasive Surgery. Adv. Intell. Syst. 2, 1900086 (2020).\n2. Piskarev, Y. et al. Fast-Response Variable-Stiffness Magnetic Catheters for Minimally Invasive Surgery. Adv. Sci. 11, 2305537 (2024).\n3. Kim, J., Choi, W. Y., Kang, S., Kim, C. & Cho, K. J. Continuously Variable Stiffness Mechanism Using Nonuniform Patterns on Coaxial Tubes for Continuum Microsurgical Robot. IEEE Trans. Robot. 35, 1475–1487 (2019).\n4. Mao, L. et al. Magnetic steering continuum robot for transluminal procedures with programmable shape and functionalities. Nat. Commun. 15, 3759 (2024).\n5. Lussi, J. et al. A Submillimeter Continuous Variable Stiffness Catheter for Compliance Control. Adv. Sci. 8, 2101290 (2021).\n6. Sun, Y. et al. Instant variable stiffness in cardiovascular catheters based on fiber jamming. Sci. Adv. 11, eadn1207 (2025).\n7. Yang, P., Mao, L., Tian, C., Meng, X. & Xie, H. A Cooperative and Multifunctional Magnetic Continuum Robot for Noninteractive Access, Dexterous Navigation, and Versatile Manipulation. Adv. Funct. Mater. 35, 2412543 (2024).\n8. Piskarev, Y. et al. A Variable Stiffness Magnetic Catheter Made of a Conductive Phase-Change Polymer for Minimally Invasive Surgery. Adv. Funct. Mater. 32, 202107662 (2022).\n9. Li, Z. et al. Design and Hierarchical Control of a Homocentric Variable-Stiffness Magnetic Catheter for Multiarm Robotic Ultrasound-Assisted Coronary Intervention. IEEE Trans. Robot. 40, 2306–2326 (2024).\n10. Mattmann, M. et al. Thermoset Shape Memory Polymer Variable Stiffness 4D Robotic Catheters. Adv. Sci. 9, 202103277 (2022).\n11. Liu, H. et al. A concentric tube magnetic continuum robot with multiple stiffness levels and high flexibility for potential endovascular intervention. J. Magn. Magn. Mater. 597, 172023 (2024)."}, {"round": 2, "reviewer_comments": {"reviewer_1": "The revised manuscript has addressed all my comments and improved the quality and impact of the study with additional in vitro and in vivo experiments, data, and analysis. Also, multi-spring stiffness tuning extends the scalability of the proposed method. Overall, it is a very interesting study and an important contribution to the magnetic robotic community. I have no further comments.", "reviewer_2": "I am satisfied with the substantial improvements in the revised manuscript. The added model depth, inclusion of animal studies, and overall rigor represent a marked advance over the initial submission.\nHowever, the authors have not addressed my primary comment:\n“Major comments\nLimited conceptual novelty relative to prior art on magnetic, variable-stiffness catheters\nThe manuscript presents a helix-sliding sleeve that shortens/elongates a soft segment to achieve a 40-fold stiffness change, coupled with radial-magnet torque steering. Yet the key ingredients (i) a magnetically actuated rigid helix, (ii) sliding-sleeve stiffness modulation, and (iii) decoupled torque steering each appear in the literature individually or in combination. Representative examples include LMPA-based or SMP-based magnetic catheters with multi-segment stiffness switching (Chautems et al., 2020; Piskarev et al., 2023) as well as continuum robots that integrate ferromagnetic steering with mechanically programmable stiffness (Kim et al., 2019; Mao et al., 2024).\nThe manuscript, therefore, risks being seen as an incremental combination rather than a substantive advance.”\nIt would be beneficial to state the central hypothesis clearly and to articulate the specific advances made by this study."}, "author_response": ""}]

Decision letter