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        Based on the interdisciplinary intersection of physics and life sciences, diagnostic and therapeutic strategies based on precision medicine have recently attracted considerable attention due to the practical applicability of new engineering methods in many fields of medicine, especially in oncology. Within this framework, the use of ultrasound to attack cancer cells in tumors in order to cause possible mechanical damage on various scales is attracting increasing attention from scientists around the world. Taking these factors into account, based on elastodynamic timing solutions and numerical simulations, we present a preliminary study of computer simulation of ultrasound propagation in tissues in order to select suitable frequencies and powers by local irradiation. New diagnostic platform for the laboratory On-Fiber technology, called the hospital needle and already patented. It is believed that the results of the analysis and related biophysical insights could pave the way for new integrated diagnostic and therapeutic approaches that could play a central role in the application of precision medicine in the future, drawing from the fields of physics. A growing synergy between biology is beginning .
        With the optimization of a large number of clinical applications, the need to reduce side effects on patients gradually began to emerge. To this end, precision medicine1, 2, 3, 4, 5 has become a strategic goal to reduce the dose of drugs delivered to patients, essentially following two main approaches. The first is based on a treatment designed according to the patient’s genomic profile. The second, which is becoming the gold standard in oncology, aims to avoid systemic drug delivery procedures by trying to release a small amount of drug, while at the same time increasing accuracy through the use of local therapy. The ultimate goal is to eliminate or at least minimize the negative effects of many therapeutic approaches, such as chemotherapy or systemic administration of radionuclides. Depending on the type of cancer, location, radiation dose, and other factors, even radiation therapy can have a high inherent risk to healthy tissue. In the treatment of glioblastoma6,7,8,9 surgery successfully removes the underlying cancer, but even in the absence of metastases, many small cancerous infiltrates may be present. If they are not completely removed, new cancerous masses can grow within a relatively short period of time. In this context, the aforementioned precision medicine strategies are difficult to apply because these infiltrates are difficult to detect and spread over a large area. These barriers prevent definitive results in preventing any recurrence with precision medicine, so systemic delivery methods are preferred in some cases, although the drugs used can have very high levels of toxicity. To overcome this problem, the ideal treatment approach would be to use minimally invasive strategies that can selectively attack cancer cells without affecting healthy tissue. In light of this argument, the use of ultrasonic vibrations, which have been shown to affect cancerous and healthy cells differently, both in unicellular systems and in mesoscale heterogeneous clusters, seems like a possible solution.
        From a mechanistic point of view, healthy and cancerous cells actually have different natural resonant frequencies. This property is associated with oncogenic changes in the mechanical properties of the cytoskeletal structure of cancer cells12,13, while tumor cells are, on average, more deformable than normal cells. Thus, with an optimal choice of ultrasound frequency for stimulation, vibrations induced in selected areas can cause damage to living cancerous structures, minimizing the impact on the healthy environment of the host. These not yet fully understood effects may include destruction of certain cellular structural components due to high-frequency vibrations induced by ultrasound (in principle very similar to lithotripsy14) and cellular damage due to a phenomenon similar to mechanical fatigue, which in turn can change cellular structure. programming and mechanobiology. Although this theoretical solution seems to be very suitable, unfortunately it cannot be used in cases where anechoic biological structures prevent the direct application of ultrasound, for example, in intracranial applications due to the presence of bone, and some breast tumor masses are located in adipose tissue. Attenuation may limit the site of potential therapeutic effect. To overcome these problems, ultrasound must be applied locally with specially designed transducers that can reach the irradiated site as less invasively as possible. With this in mind, we considered the possibility of using ideas related to the possibility of creating an innovative technological platform called the “needle hospital”15. The “Hospital in the Needle” concept involves the development of a minimally invasive medical instrument for diagnostic and therapeutic applications, based on the combination of various functions in one medical needle. As discussed in more detail in the Hospital Needle section, this compact device is primarily based on the advantages of 16, 17, 18, 19, 20, 21 fiber optic probes, which, due to their characteristics, are suitable for insertion into standard 20 medical needles, 22 lumens. Leveraging the flexibility afforded by Lab-on-Fiber (LOF)23 technology, fiber is effectively becoming a unique platform for miniaturized and ready-to-use diagnostic and therapeutic devices, including fluid biopsy and tissue biopsy devices. in biomolecular detection24,25, light-guided local drug delivery26,27, high-precision local ultrasound imaging28, thermal therapy29,30 and spectroscopy-based cancer tissue identification31. Within this concept, using a localization approach based on the “needle in the hospital” device, we investigate the possibility of optimizing local stimulation of resident biological structures by using the propagation of ultrasound waves through needles to excite ultrasound waves within the region of interest. . Thus, low-intensity therapeutic ultrasound can be applied directly to the risk area with minimal invasiveness for sonicating cells and small solid formations in soft tissues, as in the case of the aforementioned intracranial surgery, a small hole in the skull must be inserted with a needle. Inspired by recent theoretical and experimental results suggesting that ultrasound can halt or delay the development of certain cancers,32,33,34 the proposed approach may help address, at least in principle, the key trade-offs between aggressive and curative effects. With these considerations in mind, in the present paper, we investigate the possibility of using an in-hospital needle device for minimally invasive ultrasound therapy for cancer. More precisely, in the Scattering Analysis of Spherical Tumor Masses for Estimating Growth-Dependent Ultrasound Frequency section, we use well-established elastodynamic methods and acoustic scattering theory to predict the size of spherical solid tumors grown in an elastic medium. stiffness that occurs between the tumor and host tissue due to growth-induced remodeling of the material. Having described our system, which we call the “Hospital in the Needle” section, in the “Hospital in the Needle” section, we analyze the propagation of ultrasonic waves through medical needles at the predicted frequencies and their numerical model irradiates the environment to study the main geometric parameters (the actual inner diameter , length and sharpness of the needle), affecting the transmission of the acoustic power of the instrument. Given the need to develop new engineering strategies for precision medicine, it is believed that the proposed study could help develop a new tool for cancer treatment based on the use of ultrasound delivered through an integrated theragnostic platform that integrates ultrasound with other solutions. Combined, such as targeted drug delivery and real-time diagnostics within a single needle.
        The effectiveness of providing mechanistic strategies for the treatment of localized solid tumors using ultrasonic (ultrasound) stimulation has been the goal of several papers dealing both theoretically and experimentally with the effect of low-intensity ultrasonic vibrations on single-cell systems 10, 11, 12. , 32, 33, 34, 35, 36 Using viscoelastic models, several investigators have analytically demonstrated that tumor and healthy cells exhibit different frequency responses characterized by distinct resonant peaks in the US 10,11,12 range. This result suggests that, in principle, tumor cells can be selectively attacked by mechanical stimuli that preserve the host environment. This behavior is a direct consequence of key evidence that, in most cases, tumor cells are more malleable than healthy cells, possibly to enhance their ability to proliferate and migrate37,38,39,40. Based on the results obtained with single cell models, eg at the microscale, the selectivity of cancer cells has also been demonstrated at the mesoscale through numerical studies of the harmonic responses of heterogeneous cell aggregates. Providing a different percentage of cancer cells and healthy cells, multicellular aggregates hundreds of micrometers in size were built hierarchically. At the mesolevel of these aggregates, some microscopic features of interest are preserved due to the direct implementation of the main structural elements that characterize the mechanical behavior of single cells. In particular, each cell uses a tensegrity-based architecture to mimic the response of various prestressed cytoskeletal structures, thereby affecting their overall stiffness12,13. Theoretical predictions and in vitro experiments of the above literature have given encouraging results, indicating the need to study the sensitivity of tumor masses to low-intensity therapeutic ultrasound (LITUS), and the assessment of the frequency of irradiation of tumor masses is crucial. position LITUS for on-site application.
        However, at the tissue level, the submacroscopic description of the individual component is inevitably lost, and the properties of the tumor tissue can be traced using sequential methods to track the mass growth and stress-induced remodeling processes, taking into account the macroscopic effects of growth. -induced changes in tissue elasticity on a scale of 41.42. Indeed, unlike unicellular and aggregate systems, solid tumor masses grow in soft tissues due to the gradual accumulation of aberrant residual stresses, which change the natural mechanical properties due to an increase in the overall intratumoral rigidity, and tumor sclerosis often becomes a determining factor in tumor detection.
        With these considerations in mind, here we analyze the sonodynamic response of tumor spheroids modeled as elastic spherical inclusions growing in a normal tissue environment. More precisely, the elastic properties associated with the stage of the tumor were determined based on the theoretical and experimental results obtained by some authors in previous work. Among them, the evolution of solid tumor spheroids grown in vivo in heterogeneous media has been studied by applying non-linear mechanical models 41,43,44 in combination with interspecies dynamics to predict the development of tumor masses and associated intratumoral stress. As mentioned above, growth (eg, inelastic prestretching) and residual stress cause progressive remodeling of the properties of the tumor material, thereby also changing its acoustic response. It is important to note that in ref. 41 the co-evolution of growth and solid stress in tumors has been demonstrated in experimental campaigns in animal models. In particular, a comparison of the stiffness of breast tumor masses resected at different stages with the stiffness obtained by reproducing similar conditions in silico on a spherical finite element model with the same dimensions and taking into account the predicted residual stress field confirmed the proposed method of model validity. . In this work, previously obtained theoretical and experimental results are used to develop a new developed therapeutic strategy. In particular, predicted sizes with corresponding evolutionary resistance properties were calculated here, which were thus used to estimate the frequency ranges to which tumor masses embedded in the host environment are more sensitive. To this end, we thus investigated the dynamic behavior of the tumor mass at different stages, taken at different stages, taking into account acoustic indicators in accordance with the generally accepted principle of scattering in response to ultrasonic stimuli and highlighting possible resonant phenomena of the spheroid. depending on tumor and host Growth-dependent differences in stiffness between tissues.
        Thus, tumor masses were modeled as elastic spheres of radius \(a\) in the surrounding elastic environment of the host based on experimental data showing how bulky malignant structures grow in situ in spherical shapes. Referring to Figure 1, using the spherical coordinates \(\{ r,\theta ,\varphi \}\) (where \(\theta\) and \(\varphi\) represent the anomaly angle and azimuth angle respectively), the tumor domain occupies Region embedded in healthy space \({\mathcal {V}}_{T}=\{ (r,\theta ,\varphi ):r\le a\}\) unbounded region \({\mathcal {V} }_{H} = \{ (r,\theta,\varphi):r > a\}\). Referring to Supplementary Information (SI) for a complete description of the mathematical model based on the well-established elastodynamic basis reported in many literatures45,46,47,48, we consider here a problem characterized by an axisymmetric oscillation mode. This assumption implies that all variables within the tumor and healthy areas are independent of the azimuthal coordinate \(\varphi\) and that no distortion occurs in this direction. Consequently, the displacement and stress fields can be obtained from two scalar potentials \(\phi = \hat{\phi}\left( {r,\theta} \right)e^{{ – i \omega {\kern 1pt } t }}\) and \(\chi = \hat{\chi }\left( {r,\theta } \right)e^{{ – i\omega {\kern 1pt} t }}\) , they are respectively related with a longitudinal wave and a shear wave, the coincidence time t between the surge \(\theta \) and the angle between the direction of the incident wave and the position vector \({\mathbf {x))\) (as shown in figure 1) and \(\omega = 2\pi f\) represents the angular frequency. In particular, the incident field is modeled by the plane wave \(\phi_{H}^{(in)}\) (also introduced in the SI system, in equation (A.9)) propagating into the volume of the body according to the law expression
        where \(\phi_{0}\) is the amplitude parameter. The spherical expansion of an incident plane wave (1) using a spherical wave function is the standard argument:
        Where \(j_{n}\) is the spherical Bessel function of the first kind of order \(n\), and \(P_{n}\) is the Legendre polynomial. Part of the incident wave of the investment sphere is scattered in the surrounding medium and overlaps the incident field, while the other part is scattered inside the sphere, contributing to its vibration. To do this, the harmonic solutions of the wave equation \(\nabla^{2} \hat{\phi } + k_{1}^{2} {\mkern 1mu} \hat{\phi } = 0\,\ ) and \ (\ nabla^{2} {\mkern 1mu} \hat{\chi } + k_{2}^{2} \hat{\chi } = 0\), provided for example by Eringen45 (see also SI ) may indicate tumor and healthy areas. In particular, scattered expansion waves and isovolumic waves generated in the host medium \(H\) admit their respective potential energies:
        Among them, the spherical Hankel function of the first kind \(h_{n}^{(1)}\) is used to consider the outgoing scattered wave, and \(\alpha_{n}\) and \(\beta_{ n}\ ) are the unknowns coefficients. in the equation. In equations (2)–(4), the terms \(k_{H1}\) and \(k_{H2}\) denote the wave numbers of rarefaction and transverse waves in the main area of ​​the body, respectively (see SI). Compression fields inside the tumor and shifts have the form
        Where \(k_{T1}\) and \(k_{T2}\) represent the longitudinal and transverse wave numbers in the tumor region, and the unknown coefficients are \(\gamma_{n} {\mkern 1mu}\) , \(\ eta_{n} {\mkern 1mu}\). Based on these results, non-zero radial and circumferential displacement components are characteristic of healthy regions in the problem under consideration, such as \(u_{Hr}\) and \(u_{H\theta}\) (\(u_{ H\ varphi }\ ) the symmetry assumption is no longer needed) — can be obtained from the relation \(u_{Hr} = \partial_{r} \left( {\phi + \partial_{r} (r\chi )} \right) + k_}^{2 } {\mkern 1mu} r\chi\) and \(u_{H\theta} = r^{- 1} \partial_{\theta} \left({\phi + \partial_{r } ( r\chi ) } \right)\) by forming \(\phi = \phi_{H}^{(in)} + \phi_{H}^{(s)}\) and \(\chi = \chi_ {H}^ {(s)}\) (see SI for detailed mathematical derivation). Similarly, replacing \(\phi = \phi_{T}^{(s)}\) and \(\chi = \chi_{T}^{(s)}\) returns {Tr} = \partial_{r} \left( {\phi + \partial_{r} (r\chi)} \right) + k_{T2}^{2} {\mkern 1mu} r\chi\) and \(u_{T\theta} = r^{-1}\partial _{\theta }\left({\phi +\partial_{r}(r\chi )}\right)\).
       (Left) Geometry of a spherical tumor grown in a healthy environment through which an incident field propagates, (right) Corresponding evolution of the tumor-host stiffness ratio as a function of tumor radius, reported data (adapted from Carotenuto et al. 41) from in compression tests vitro were obtained from solid breast tumors inoculated with MDA-MB-231 cells.
        Assuming linear elastic and isotropic materials, the non-zero stress components in the healthy and tumor regions, i.e. \(\sigma_{Hpq}\) and \(\sigma_{Tpq}\) – obey the generalized Hooke’s law, given that there are different Lamé moduli , which characterize host and tumor elasticity, denoted as \(\{ \mu_{H},\,\lambda_{H} \}\) and \(\{ \mu_{T},\, \lambda_{T} \ }\) (see Equation (A.11) for the full expression of the stress components represented in SI). In particular, according to the data in reference 41 and presented in Figure 1, growing tumors showed a change in tissue elasticity constants. Thus, displacements and stresses in the host and tumor regions are determined completely up to a set of unknown constants \({{ \varvec{\upxi}}}_{n} = \{ \alpha_{n} ,{\mkern 1mu } \ beta_{ n} {\mkern 1mu} \gamma_{n} ,\eta_{n} \}\ ) has theoretically infinite dimensions. To find these coefficient vectors, suitable interfaces and boundary conditions between the tumor and healthy areas are introduced. Assuming perfect binding at the tumor-host interface \(r = a\), continuity of displacements and stresses requires the following conditions:
        System (7) forms a system of equations with infinite solutions. In addition, each boundary condition will depend on the anomaly \(\theta\). To reduce the boundary value problem to a complete algebraic problem with \(N\) sets of closed systems, each of which is in the unknown \({{\varvec{\upxi}}}_{n} = \{ \alpha_{n},{ \mkern 1mu} \beta_{n} {\mkern 1mu} \gamma_{n}, \eta_{n} \}_{n = 0,…,N}\) (with \ ( N \to \infty \), theoretically), and to eliminate the dependence of the equations on the trigonometric terms, the interface conditions are written in a weak form using the orthogonality of the Legendre polynomials. In particular, the equation (7)1,2 and (7)3,4 are multiplied by \(P_{n} \left( {\cos \theta} \right)\) and \(P_{n}^{1} \left( { \cos\theta}\right)\) and then integrate between \(0\) and \(\pi\) using mathematical identities:
       Thus, the interface condition (7) returns a quadratic algebraic equation system, which can be expressed in matrix form as \({\mathbb{D}}_{n} (a) \cdot {{\varvec{\upxi }}} _{ n} = {\mathbf{q}}_{n} (a)\) and get the unknown \({{\varvec{\upxi}}}_{n}\ ) by solving Cramer’s rule .
        To estimate the energy flux scattered by the sphere and obtain information about its acoustic response based on data on the scattered field propagating in the host medium, an acoustic quantity is of interest, which is a normalized bistatic scattering cross section. In particular, the scattering cross section, denoted \(s), expresses the ratio between the acoustic power transmitted by the scattered signal and the division of energy carried by the incident wave. In this regard, the magnitude of the shape function \(\left| {F_{\infty} \left(\theta \right)} \right|^{2}\) is a frequently used quantity in the study of acoustic mechanisms embedded in a liquid or solid Scattering of objects in the sediment. More precisely, the amplitude of the shape function is defined as the differential scattering cross section \(ds\) per unit area, which differs by the normal to the direction of propagation of the incident wave:
       where \(f_{n}^{pp}\) and \(f_{n}^{ps}\) denote the modal function, which refers to the ratio of the powers of the longitudinal wave and the scattered wave relative to the incident P-wave in the receiving medium, respectively, are given with the following expressions:
        Partial wave functions (10) can be studied independently in accordance with the resonant scattering theory (RST)49,50,51,52, which makes it possible to separate the target elasticity from the total stray field when studying different modes. According to this method, the modal form function can be decomposed into a sum of two equal parts, namely \(f_{n} = f_{n}^{(res)} + f_{n}^{(b)}\ ) are related to the resonant and nonresonant background amplitudes, respectively. The shape function of the resonant mode is related to the response of the target, while the background is usually related to the shape of the scatterer. To detect the first formant of the target for each mode, the amplitude of the modal resonance shape function \(\left| {f_{n}^{(res)} \left( \theta \right)} \right|\ ) is calculated assuming a hard background, consisting of impenetrable spheres in an elastic host material. This hypothesis is motivated by the fact that, in general, both stiffness and density increase with the growth of the tumor mass due to the residual compressive stress. Thus, at a severe level of growth, the impedance ratio \(\rho_{T} c_{1T} /\rho_{H} c_{1H}\) is expected to be greater than 1 for most macroscopic solid tumors developing in soft tissues. For example, Krouskop et al. 53 reported a ratio of cancerous to normal modulus of about 4 for prostate tissue, while this value increased to 20 for breast tissue samples. These relationships inevitably change the acoustic impedance of the tissue, as also demonstrated by elastography analysis54,55,56, and may be related to localized tissue thickening caused by tumor hyperproliferation. This difference has also been observed experimentally with simple compression tests of breast tumor blocks grown at different stages32, and remodeling of the material can be well followed with predictive cross-species models of non-linearly growing tumors43,44. The stiffness data obtained are directly related to the evolution of the Young’s modulus of solid tumors according to the formula \(E_{T} = S\left( {1 – \nu ^{2} } \right)/a\sqrt \varepsilon\ )( spheres with radius \(a\), stiffness \(S\) and Poisson’s ratio \(\nu\) between two rigid plates 57, as shown in Figure 1). Thus, it is possible to obtain acoustic impedance measurements of the tumor and the host at different growth levels. In particular, in comparison with the modulus of normal tissue equal to 2 kPa in Fig. 1, the elastic modulus of breast tumors in the volume range of about 500 to 1250 mm3 resulted in an increase from about 10 kPa to 16 kPa, which is consistent with the reported data. in references 58, 59 it was found that the pressure in breast tissue samples is 0.25–4 kPa with vanishing precompression. Also assume that the Poisson’s ratio of an almost incompressible tissue is 41.60, which means that the density of the tissue does not change significantly as the volume increases. In particular, the average mass population density \(\rho = 945\,{\text{kg}}\,{\text{m}}^{ – 3}\)61 is used. With these considerations, stiffness can take on a background mode using the following expression:
        Where the unknown constant \(\widehat{{{\varvec{\upxi))))_{n} = \{\delta_{n} ,\upsilon_{n} \}\) can be calculated taking into account the continuity bias ( 7 )2,4, that is, by solving the algebraic system \(\widehat{{\mathbb{D}}}_{n} (a) \cdot \widehat{({\varvec{\upxi}}} } _{n } = \widehat{{\mathbf{q}}}_{n} (a)\) involving minors\(\widehat{{\mathbb{D}}}_{n} (a) = \ { { \ mathbb{D}}_{n} (a)\}_{{\{ (1,3),(1,3)\} }}\) and the corresponding simplified column vector\(\widehat{ {\mathbf {q}}}_{n} (а)\). Provides basic knowledge in equation. (11), two amplitudes of the backscattering resonant mode function \(\left| {f_{n}^{{\left( {res} \right)\,pp}} \left( \theta \right)} \right| = \left|{f_{n}^{pp} \left( \theta \right) – f_{n}^{pp(b)} \left( \theta \right)} \right|\) and \( \left|{f_{n}^{{\left( {res} \right)\,ps}} \left( \theta \right)} \right|= \left|{f_{n}^{ps} \left( \theta \right) – f_{n}^{ps(b)} \left( \theta \right)} \right|\) refers to P-wave excitation and P- and S-wave reflection, respectively. Further, the first amplitude was estimated as \(\theta = \pi\), and the second amplitude was estimated as \(\theta = \pi/4\). By loading various composition properties. Figure 2 shows that the resonant features of tumor spheroids up to about 15 mm in diameter are mainly concentrated in the frequency band of 50-400 kHz, which indicates the possibility of using low-frequency ultrasound to induce resonant tumor excitation. cells. A lot of. In this frequency band, the RST analysis revealed single-mode formants for modes 1 to 6, highlighted in Figure 3. Here, both pp- and ps-scattered waves show formants of the first type, occurring at very low frequencies, which increase from about 20 kHz for the mode 1 to about 60 kHz for n = 6, showing no significant difference in sphere radius. The resonant function ps then decays, while the combination of large amplitude pp formants provides a periodicity of about 60 kHz, showing a higher frequency shift with increasing mode number. All analyzes were performed using Mathematica®62 computing software.
        The backscatter form functions obtained from the module of breast tumors of different sizes are shown in Fig. 1, where the highest scattering bands are highlighted taking into account mode superposition.
       Resonances of selected modes from \(n = 1\) to \(n = 6\), calculated upon excitation and reflection of the P-wave at different tumor sizes (black curves from \(\left | {f_{ n} ^{{\ left( {res} \right)\,pp}} \left( \pi \right)} \right| = \left| {f_{n}^{pp} \left ( \pi \ right) – f_{n }^{pp(b)} \left( \pi \right)} \right|\)) and P-wave excitation and S-wave reflection (gray curves given by modal shape function \( \left | {f_{n }^{{\left( {res} \right)\,ps}} \left( {\pi /4} \right)} \right| = \left| {f_{n} ^{ ps} \left( {\pi /4} \right) – f_{n}^{ps(b)} \left( {\pi /4} \right)} \right |\)).
        The results of this preliminary analysis using far-field propagation conditions can guide the selection of drive-specific drive frequencies in the following numerical simulations to study the effect of microvibration stress on mass. The results show that the calibration of optimal frequencies can be stage-specific during tumor growth and can be determined using the results of growth models to establish biomechanical strategies used in disease therapy to correctly predict tissue remodeling.
        Significant advances in nanotechnology are driving the scientific community to find new solutions and methods to develop miniaturized and minimally invasive medical devices for in vivo applications. In this context, LOF technology has shown a remarkable ability to expand the capabilities of optical fibers, enabling the development of new minimally invasive fiber optic devices for life science applications21, 63, 64, 65. The idea of ​​integrating 2D and 3D materials with desired chemical, biological, and optical properties on the sides 25 and/or ends 64 of optical fibers with full spatial control at the nanoscale leads to the emergence of a new class of fiber optic nanooptodes. has a wide range of diagnostic and therapeutic functions. Interestingly, due to their geometric and mechanical properties (small cross section, large aspect ratio, flexibility, low weight) and the biocompatibility of materials (usually glass or polymers), optical fibers are well suited for insertion into needles and catheters. Medical applications20, paving the way for a new vision of the “needle hospital” (see Figure 4).
        In fact, due to the degrees of freedom afforded by LOF technology, by utilizing the integration of micro- and nanostructures made from various metallic and/or dielectric materials, optical fibers can be properly functionalized for specific applications often supporting resonant mode excitation. , The light field 21 is strongly positioned. The containment of light on a subwavelength scale, often in combination with chemical and/or biological processing63 and the integration of sensitive materials such as smart polymers65,66 can enhance control over the interaction of light and matter, which can be useful for theranostic purposes. The choice of type and size of integrated components/materials obviously depends on the physical, biological or chemical parameters to be detected21,63.
        Integration of LOF probes into medical needles directed to specific sites in the body will enable local fluid and tissue biopsies in vivo, allowing simultaneous local treatment, reducing side effects and increasing efficiency. Potential opportunities include the detection of various circulating biomolecules, including cancer. biomarkers or microRNAs (miRNAs)67, identification of cancerous tissues using linear and non-linear spectroscopy such as Raman spectroscopy (SERS)31, high-resolution photoacoustic imaging22,28,68, laser surgery and ablation69, and local delivery drugs using light27 and automatic guidance of needles into the human body20. It is worth noting that although the use of optical fibers avoids the typical disadvantages of “classical” methods based on electronic components, such as the need for electrical connections and the presence of electromagnetic interference, this allows various LOF sensors to be effectively integrated into the system. single medical needle. Particular attention must be paid to reducing harmful effects such as pollution, optical interference, physical obstructions that cause crosstalk effects between different functions. However, it is also true that many of the functions mentioned do not have to be active at the same time. This aspect makes it possible to at least reduce interference, thereby limiting the negative impact on the performance of each probe and the accuracy of the procedure. These considerations allow us to view the concept of the “needle in the hospital” as a simple vision to lay a solid foundation for the next generation of therapeutic needles in the life sciences.
       With regard to the specific application discussed in this paper, in the next section we will numerically investigate the ability of a medical needle to direct ultrasonic waves into human tissues using their propagation along its axis.
       Propagation of ultrasonic waves through a medical needle filled with water and inserted into soft tissues (see diagram in Fig. 5a) was modeled using the commercial Comsol Multiphysics software based on the finite element method (FEM)70, where the needle and tissue are modeled as linear elastic environment.
        Referring to Figure 5b, the needle is modeled as a hollow cylinder (also known as a “cannula”) made of stainless steel, a standard material for medical needles71. In particular, it was modeled with Young’s modulus E = 205 GPa, Poisson’s ratio ν = 0.28, and density ρ = 7850 kg m −372.73. Geometrically, the needle is characterized by a length L, an internal diameter D (also called “clearance”) and a wall thickness t. In addition, the tip of the needle is considered to be inclined at an angle α with respect to the longitudinal direction (z). The volume of water essentially corresponds to the shape of the inner region of the needle. In this preliminary analysis, the needle was assumed to be completely immersed in a region of tissue (assumed to extend indefinitely), modeled as a sphere of radius rs, which remained constant at 85 mm during all simulations. In more detail, we finish the spherical region with a perfectly matched layer (PML), which at least reduces unwanted waves reflected from “imaginary” boundaries. We then chose the radius rs so as to place the spherical domain boundary far enough from the needle not to affect the computational solution, and small enough not to affect the computational cost of the simulation.
        A harmonic longitudinal shift of frequency f and amplitude A is applied to the lower boundary of the stylus geometry; this situation represents an input stimulus applied to the simulated geometry. At the remaining boundaries of the needle (in contact with tissue and water), the accepted model is considered to include a relationship between two physical phenomena, one of which is related to structural mechanics (for the area of ​​the needle), and the other to structural mechanics. (for the acicular region), so the corresponding conditions are imposed on the acoustics (for water and the acicular region)74. In particular, small vibrations applied to the needle seat cause small voltage perturbations; thus, assuming that the needle behaves like an elastic medium, the displacement vector U can be estimated from the elastodynamic equilibrium equation (Navier)75. Structural oscillations of the needle cause changes in the water pressure inside it (considered to be stationary in our model), as a result of which sound waves propagate in the longitudinal direction of the needle, essentially obeying the Helmholtz equation76. Finally, assuming that the nonlinear effects in tissues are negligible and that the amplitude of the shear waves is much smaller than the amplitude of the pressure waves, the Helmholtz equation can also be used to model the propagation of acoustic waves in soft tissues. After this approximation, the tissue is considered as a liquid77 with a density of 1000 kg/m3 and a speed of sound of 1540 m/s (ignoring frequency-dependent damping effects). To connect these two physical fields, it is necessary to ensure the continuity of normal movement at the boundary of the solid and liquid, the static equilibrium between pressure and stress perpendicular to the boundary of the solid, and the tangential stress at the boundary of the liquid must be equal to zero. 75 .
        In our analysis, we investigate the propagation of acoustic waves along a needle under stationary conditions, focusing on the influence of the geometry of the needle on the emission of waves inside the tissue. In particular, we investigated the influence of the inner diameter of the needle D, the length L and the bevel angle α, keeping the thickness t fixed at 500 µm for all cases studied. This value of t is close to the typical standard wall thickness 71 for commercial needles.
        Without loss of generality, the frequency f of the harmonic displacement applied to the base of the needle was taken equal to 100 kHz, and the amplitude A was 1 μm. In particular, the frequency was set to 100 kHz, which is consistent with the analytical estimates given in the section “Scattering analysis of spherical tumor masses to estimate growth-dependent ultrasound frequencies”, where a resonance-like behavior of tumor masses was found in the frequency range of 50–400 kHz, with the largest scattering amplitude concentrated at lower frequencies around 100–200 kHz (see Fig. 2).
        The first parameter studied was the internal diameter D of the needle. For convenience, it is defined as an integer fraction of the acoustic wave length in the cavity of the needle (i.e., in water λW = 1.5 mm). Indeed, the phenomena of wave propagation in devices characterized by a given geometry (for example, in a waveguide) often depend on the characteristic size of the geometry used in comparison with the wavelength of the propagating wave. In addition, in the first analysis, in order to better emphasize the effect of the diameter D on the propagation of the acoustic wave through the needle, we considered a flat tip, setting the angle α = 90°. During this analysis, the needle length L was fixed at 70 mm.
        On fig. 6a shows the average sound intensity as a function of the dimensionless scale parameter SD, i.e. D = λW/SD evaluated in a sphere with a radius of 10 mm centered on the corresponding needle tip. The scaling parameter SD changes from 2 to 6, i.e. we consider D values ​​ranging from 7.5 mm to 2.5 mm (at f = 100 kHz). The range also includes a standard value of 71 for stainless steel medical needles. As expected, the inner diameter of the needle affects the intensity of the sound emitted by the needle, with a maximum value (1030 W/m2) corresponding to D = λW/3 (i.e. D = 5 mm) and a decreasing trend with decreasing diameter. It should be taken into account that the diameter D is a geometric parameter that also affects the invasiveness of a medical device, so this critical aspect cannot be ignored when choosing the optimal value. Therefore, although the decrease in D occurs due to the lower transmission of acoustic intensity in the tissues, for the following studies, the diameter D = λW/5, i.e. D = 3 mm (corresponds to the 11G71 standard at f = 100 kHz), is considered a reasonable compromise between device intrusiveness and sound intensity transmission (average about 450 W/m2).
        The average intensity of the sound emitted by the tip of the needle (considered flat), depending on the inner diameter of the needle (a), length (b) and bevel angle α (c). The length in (a, c) is 90 mm, and the diameter in (b, c) is 3 mm.
        The next parameter to be analyzed is the length of the needle L. As per the previous case study, we consider an oblique angle α = 90° and the length is scaled as a multiple of the wavelength in water, i.e. consider L = SL λW. The dimensionless scale parameter SL is changed from 3 by 7, thus estimating the average intensity of the sound emitted by the tip of the needle in the length range from 4.5 to 10.5 mm. This range includes typical values ​​for commercial needles. The results are shown in fig. 6b, showing that the length of the needle, L, has a great influence on the transmission of sound intensity in tissues. Specifically, the optimization of this parameter made it possible to improve the transmission by about an order of magnitude. In fact, in the analyzed length range, the average sound intensity takes on a local maximum of 3116 W/m2 at SL = 4 (i.e., L = 60 mm), and the other corresponds to SL = 6 (i.e., L = 90 mm).
        After analyzing the influence of the diameter and length of the needle on the propagation of ultrasound in cylindrical geometry, we focused on the influence of the bevel angle on the transmission of sound intensity in tissues. The average intensity of the sound emanating from the fiber tip was evaluated as a function of the angle α, changing its value from 10° (sharp tip) to 90° (flat tip). In this case, the radius of the integrating sphere around the considered tip of the needle was 20 mm, so that for all values ​​of α, the tip of the needle was included in the volume calculated from the average.
        As shown in fig. 6c, when the tip is sharpened, i.e., when α decreases starting from 90°, the intensity of the transmitted sound increases, reaching a maximum value of about 1.5 × 105 W/m2, which corresponds to α = 50°, i.e. i.e., 2 is an order of magnitude higher relative to the flat state. With further sharpening of the tip (i.e., at α below 50°), the sound intensity tends to decrease, reaching values ​​comparable to a flattened tip. However, although we considered a wide range of bevel angles for our simulations, it is worth considering that sharpening the tip is necessary to facilitate insertion of the needle into the tissue. In fact, a smaller bevel angle (about 10°) can reduce the force 78 required to penetrate tissue.
        In addition to the value of the sound intensity transmitted within the tissue, the bevel angle also affects the direction of wave propagation, as shown in the sound pressure level graphs shown in Fig. 7a (for the flat tip) and 3b (for 10°). beveled tip), parallel The longitudinal direction is evaluated in the plane of symmetry (yz, cf. Fig. 5). At the extremes of these two considerations, the sound pressure level (referred to as 1 µPa) is mainly concentrated within the needle cavity (i.e. in the water) and radiated into the tissue. In more detail, in the case of a flat tip (Fig. 7a), the distribution of the sound pressure level is perfectly symmetrical with respect to the longitudinal direction, and standing waves can be distinguished in the water filling the body. The wave is oriented longitudinally (z-axis), the amplitude reaches its maximum value in water (about 240 dB) and decreases transversely, which leads to an attenuation of about 20 dB at a distance of 10 mm from the center of the needle. As expected, the introduction of a pointed tip (Fig. 7b) breaks this symmetry, and the antinodes of the standing waves “deflect” according to the tip of the needle. Apparently, this asymmetry affects the radiation intensity of the needle tip, as described earlier (Fig. 6c). To better understand this aspect, the acoustic intensity was evaluated along a cut line orthogonal to the longitudinal direction of the needle, which was located in the plane of symmetry of the needle and located at a distance of 10 mm from the tip of the needle (results in Figure 7c). More specifically, sound intensity distributions assessed at 10°, 20° and 30° oblique angles (blue, red and green solid lines, respectively) were compared to the distribution near the flat end (black dotted curves). The intensity distribution associated with flat-tipped needles appears to be symmetrical about the center of the needle. In particular, it takes on a value of about 1420 W/m2 at the center, an overflow of about 300 W/m2 at a distance of ~8 mm, and then decreases to a value of about 170 W/m2 at ~30 mm. As the tip becomes pointed, the central lobe divides into more lobes of varying intensity. More specifically, when α was 30°, three petals could be clearly distinguished in the profile measured at 1 mm from the tip of the needle. The central one is almost in the center of the needle and has an estimated value of 1850 W / m2, and the higher one on the right is about 19 mm from the center and reaches 2625 W / m2. At α = 20°, there are 2 main lobes: one per −12 mm at 1785 W/m2 and one per 14 mm at 1524 W/m2. When the tip becomes sharper and the angle reaches 10°, a maximum of 817 W/m2 is reached at about -20 mm, and three more lobes of slightly lesser intensity are visible along the profile.
        Sound pressure level in the plane of symmetry y–z of a needle with a flat end (a) and a 10° bevel (b). (c) Acoustic intensity distribution estimated along a cut line perpendicular to the longitudinal direction of the needle, at a distance of 10 mm from the tip of the needle and lying in the plane of symmetry yz. The length L is 70 mm and the diameter D is 3 mm.
        Taken together, these results demonstrate that medical needles can be effectively used to transmit ultrasound at 100 kHz into soft tissue. The intensity of the emitted sound depends on the geometry of the needle and can be optimized (subject to the limitations imposed by the invasiveness of the end device) up to values ​​in the range of 1000 W/m2 (at 10 mm). applied to the bottom of the needle 1. In the case of a micrometer offset, the needle is considered to be fully inserted into the infinitely extending soft tissue. In particular, the bevel angle strongly affects the intensity and direction of propagation of sound waves in the tissue, which primarily leads to the orthogonality of the cut of the needle tip.
        To support the development of new tumor treatment strategies based on the use of non-invasive medical techniques, the propagation of low-frequency ultrasound in the tumor environment was analyzed analytically and computationally. In particular, in the first part of the study, a temporary elastodynamic solution allowed us to study the scattering of ultrasonic waves in solid tumor spheroids of known size and stiffness in order to study the frequency sensitivity of the mass. Then, frequencies of the order of hundreds of kilohertz were chosen, and the local application of vibration stress in the tumor environment using a medical needle drive was modeled in numerical simulation by studying the influence of the main design parameters that determine the transfer of the acoustic power of the instrument to the environment. The results show that medical needles can be effectively used to irradiate tissues with ultrasound, and its intensity is closely related to the geometrical parameter of the needle, called the working acoustic wavelength. In fact, the intensity of irradiation through the tissue increases with increasing internal diameter of the needle, reaching a maximum when the diameter is three times the wavelength. The length of the needle also provides some degree of freedom to optimize exposure. The latter result is indeed maximized when the needle length is set to a certain multiple of the operating wavelength (specifically 4 and 6). Interestingly, for the frequency range of interest, the optimized diameter and length values ​​are close to those commonly used for standard commercial needles. The bevel angle, which determines the sharpness of the needle, also affects the emissivity, peaking at about 50° and providing good performance at about 10°, which is commonly used for commercial needles. . Simulation results will be used to guide the implementation and optimization of the hospital’s intraneedle diagnostic platform, integrating diagnostic and therapeutic ultrasound with other in-device therapeutic solutions and realizing collaborative precision medicine interventions.
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