WO2020014277A1

WO2020014277A1 - In vivo three-dimensional printing of biocompatible materials

Info

Publication number: WO2020014277A1
Application number: PCT/US2019/041075
Authority: WO
Inventors: Cynthia M. Furse; Huanan Zhang
Original assignee: University Of Utah Research Foundation
Priority date: 2018-07-10
Filing date: 2019-07-09
Publication date: 2020-01-16

Abstract

Disclosed are methods for in vivo 3D printing of polymer structures at a targeted anatomical location. A method includes injecting a biocompatible polymer at the targeted location, and applying focused and controlled electromagnetic hyperthermia to the polymer to induce polymerization in a spatially and temporally controlled manner.

Description

IN VIVO THREE-DIMENSIONAL PRINTING OF

BIOCOMPATIBLE MATERIALS

CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority to and the benefit of United States Provisional

Patent Application No. 62/696,290, filed July 10, 2018, and titled“Three-Dimensional Printing Within Tissue,” the entirety of which is incorporated herein by this reference.

BACKGROUND

[0002] 3D printing has utilized in engineering design, prototype development, manufacturing, and medicine. 3D printing technologies have been used to construct everything from entire houses to the smallest mechanical components. The ability to create custom 3D printed structures and components from a computer file has the potential to further revolutionize many different areas.

[0003] The medical field is an area that has generated a great deal of interest in 3D printing. When combined with advanced imaging techniques, medical implants such as artificial joints, coronary stents, and other implantable electronics can be custom printed to exactly fit the individual patient. The ability to customize medical devices and components for individual patients holds tremendous potential for improving future care.

[0004] 3D printing with plastics or metals using thermal extrusion, laser sintering, inkjet printing, and photopolymerization is becoming mainstream. Research is enabling printing of biological components, synthetic tissue matrices, and artificial organs. Conventionally, this printing is done outside the body, and the structures are then surgically implanted.

[0005] The use of 3D printing in the medical field is a new and quickly growing technology. While useful, there are several areas in which the technology remains limited. For example, as mentioned above, implantable medical devices must be formed outside of the body and then invasively implanted in the conventional sense.

[0006] There is also a medical need for shape-specific soft tissue implants and fillers. For example, major facial reconstruction surgery may use 3D printed bone replacements to repair the general bone structure of the face. However, the soft tissue surrounding the bone or bone replacement, critical for both functional support and appearance, must heal on its own, and often only coarsely remodels to its original shape. [0007] The subject matter claimed herein is not limited to embodiments that solve any disadvantages or that operate only in environments such as those described above. Rather, this background is only provided to illustrate one exemplary technology area where some embodiments described herein may be practiced.

SUMMARY

[0008] An embodiment disclosed herein includes a method for in vivo 3D printing. The method comprises injecting a biocompatible polymer material into a targeted tissue, the biocompatible polymer material comprising an enzymatically cross-linkable polymer and cross-linking enzymes encapsulated in thermosensitive capsules. The method then comprises applying heat to the biocompatible polymer material, such as through focused electromagnetic hyperthermia. The heat causes the thermosensitive capsules to degrade and release the encapsulated enzymes. The released enzymes then catalyze polymerization of the enzymatically cross-linkable polymer, causing the biocompatible polymer material to cross- link and polymerize in place within the targeted tissue.

[0009] Heat may be applied to the tissue via applied radiofrequency or microwave energy, for example. Exemplary, non-limiting methods of applying heat include: (1) radiofrequency catheter ablation (RFCA); (2) interstitial microwave heating; (3) regional focused microwave heating; and (4) ultra-focused microwave heating. The applied heat need only increase the temperature of the biocompatible polymer material to about 40° C in order to initiate cross- linking.

[0010] Other known or discovered methods of controllably heating and polymerizing the biocompatible polymer material may also be utilized. By controlling the temporal and spatial application of heat to the biocompatible polymer material, the resulting polymerized structure can be 3D printed in desired size and shape in place within the targeted tissue.

[0011] In some embodiments, the biocompatible polymer material is formed as a conductive nanocomposite that includes one or more types of conductive nanoparticles. The conductive nanoparticles may be embedded in a matrix of biocompatible polymer material to provide a material that has high electrical conductivity but low mechanical stiffness. The volume ratio of conductive nanoparticles in the polymer matrix may be adjusted to balance and control electrical conductivity and mechanical properties of the nanocomposite. Conductive biocompatible polymer materials formed in this manner may then be utilized for in vivo 3D printing and/or heat-activated cross-linking as described above in relation to other embodiments.

[0012] Conductive nanocomposite materials as described above may be utilized to form effective antenna implants for use with an implantable medical device (IMD). Such antenna implants may be 3D printed in vivo in the manner described herein, and may be utilized to enable or enhance wireless telemetry of an associated IMD and/or wireless power transfer (WPT) to the IMD. Antenna implants may be formed as lines/wires, rings, or arrays including multiple antennas.

[0013] This Summary is provided to introduce a selection of concepts in a simplified form that are further described below in the Detailed Description. This Summary is not intended to identify key features or essential features of the claimed subject matter, nor is it intended to be used as an aid in determining the scope of the claimed subject matter.

[0014] Additional features and advantages will be set forth in the description which follows, and in part will be obvious from the description, or may be learned by the practice of the teachings herein. Features and advantages of the invention may be realized and obtained by means of the instruments and combinations particularly pointed out in the appended claims. Features of the present invention will become more fully apparent from the following description and appended claims, or may be learned by the practice of the invention as set forth hereinafter.

BRIEF DESCRIPTION OF THE DRAWINGS

[0015] In order to describe the manner in which the above-recited and other advantages and features can be obtained, a more particular description of the subject matter briefly described above will be rendered by reference to specific embodiments which are illustrated in the appended drawings. Understanding that these drawings depict only typical embodiments and are not therefore to be considered to be limiting in scope, embodiments will be described and explained with additional specificity and detail through the use of the accompanying drawings in which:

[0016] Figure 1 schematically illustrates an in vivo 3D printing method where a biocompatible polymer material is injected into a targeted anatomical location, and heat is controllably applied to the biocompatible polymer material to cause selective cross-linking of the polymer material where heated; [0017] Figure 2 illustrates an enzymatic polymerization system wherein the biocompatible polymer material includes polymerizable units with cross-linkable groups that cross-link upon action by the enzyme horseradish peroxidase;

[0018] Figures 3A and 3B illustrate the formation and degradation, respectively, of liposomes which may be utilized as the thermosensitive capsules;

[0019] Figure 4 illustrates a distal tip of an RFCA probe suitable for use in an in vivo 3D printing application to simultaneously extrude polymer material and heat the polymer material to induce polymerization;

[0020] Figure 5 illustrates an example of using an antenna array to steer/control heating within a region of biocompatible polymer material; and

[0021] Figures 6A and 6B illustrate a section of tissue in cross section, showing an exemplary dipole antenna implant that may be 3D printed in place within a subject;

[0022] Figure 6C illustrates another embodiment of an antenna implant including two orthogonally positioned pairs of wires; and

[0023] Figures 7A through 8B illustrate exemplary embodiments of antenna implants including ring shaped antennas.

DETAILED DESCRIPTION

Introduction

[0024] In at least one embodiment disclosed herein, 3D printing of a biocompatible polymer material is performed directly inside the body to form solid medical structures - enabling a new and useful modality for medical treatment. In at least one embodiment, this can address a specific medical need for shape-specific medical implants, including hard tissue (e.g., bone) implants, scaffolds, and support structures, but also being particularly useful in soft tissue related applications, such as soft tissue repair and/or replacement.

[0025] Embodiments described herein may be useful in, for example, facial reconstruction surgery (for victims of accident or injury, cancers, etc.) for 3D printing bone replacements to repair the major facial structures and/or for 3D printing the soft tissue surrounding the bones, critical for both functional support and appearance.

[0026] Embodiments described herein may also be utilized for short-term implants for temporary tissue support and/or assistance in the healing process. Such short-term implants are typically later removed. The only conventional, long-term option for soft-tissue support is silicone implants, long the mainstay for breast reconstructive surgery. While these implants have improved significantly over the years, they are still difficult to ideally fit the individual patient. This is particularly true when considering swelling from the original injury or invasive surgery required for placing the silicone implants.

[0027] Embodiments described herein can address some or all of these problems by 3D printing medical structures directly within the body. These medical structures can include, but are not limited to, bone repair/replacement structures, support structures, scaffolds, fillers, and soft tissue structures such as repair/replacement structures for cartilage, ligaments, and/or tendons, for example.

[0028] In at least one embodiment, surgery to plate broken bones can be replaced by a disclosed method in which the biocompatible polymer material is injected or otherwise placed along the bone and heat is then focused along its surface to solidify it as a plate. Additionally, in at least one embodiment, cartilage in an injured knee could be replaced in vivo , rather than replacing an entire knee.

[0029] Other embodiments may provide medical structures that form part of or work in conjunction with an IMD. For example, some embodiments relate to 3D printing of biocompatible, conductive antennae for use with IMDs. The antennae may beneficially provide or enhance wireless telemetry and/or WPT capabilities of the IMD. Types of IMDs that such antennae may be utilized with include, for example, retinal implants, cochlear implants, glucose monitors, lactate concentration monitors, strain sensors (e.g., for wheelchair control applications), temperature monitors, and pacemakers, though it will be understood that number of other IMDs can benefit from enhanced wireless telemetry and/or WPT capabilities.

[0030] Figure 1 illustrates an overview of the 3D printing method. The biocompatible polymer material 100 may be administered (e.g., injected) to targeted tissue 14 within a subject 10. Heat 12 may then be controllably applied to the biocompatible polymer material to cause selective cross-linking of the polymer material where heated.

[0031] As shown, the biocompatible polymer material comprises cross-linkable polymerizable units 102 and a cross-linking agent 104 encapsulated in thermosensitive capsules 106. When heat 12 is applied to the polymer material 100, the thermosensitive capsules 106 degrade, releasing the cross-linking agent 104 and allowing polymerization/cross- linking of the polymerizable units 102. [0032] Various materials may be utilized to formulate the different components of the polymer material 100, and various methods may be utilized to apply the focused hyperthermia, as explained in greater detail below.

3D Printable Biocompatible Polymer Materials

[0033] The biocompatible polymer material preferably includes an enzymatically cross- linkable polymer. For example, referring again to Figure 1, the cross-linking agent 104 may be an enzyme and the polymerizable units 102 may be enzymatically cross-linkable. Other polymerization mechanisms include chemical, direct thermal, two-component in situ , and photopolymerization, though these are much less preferred.

[0034] Chemical cross-linking involves the addition of chemical linkers to crosslink the polymer chains. Unfortunately, many chemical cross-linkers are cytotoxic. As such, the chemical-cross-linked biopolymers have to be washed extensively before an in vivo application. Moreover, the typical chemical cross-linking rate is slow (on the order of hours), making it is difficult to apply this method in real time.

[0035] Direct thermal cross-linking may be used to circumvent cytotoxicity issues.

However, many direct thermal polymerization systems have optimal temperature ranges from 140 to 200 °C, which are too high for in vivo biological applications.

[0036] Another cross-linking method is two-component in situ polymerization. Similar to an Epoxy system, two components react and cross-link upon mixing. This method can eliminate the usage of traditional chemical cross-linkers. However, conventional methods may have spatial and temporal control issues. Moreover, reaction byproducts are often cytotoxic.

[0037] Another cross-linking method is photopolymerization, which uses light to trigger cross-linking. While this method may be utilized to provide good spatial and temporal control, light penetration through tissue is limited. Accordingly, such methods are typically only practical for transdermal applications.

[0038] Enzymatic cross-linking is a subset of chemical cross-linking reactions that utilizes enzyme catalysts to catalyze the cross-linking reaction. One example is naturally-occurring blood coagulation. When the body senses an injury, within seconds, the cells releases enzymes (with ions and cofactors) that cause platelets to coagulate and stop the bleeding. Preferred embodiments of in vivo 3D printing therefore utilize enzymatic cross-linking. [0039] Figure 2 illustrates one exemplary enzymatic polymerization system. In the example of Figure 2, the polymerizable units 102 include cross-linkable groups 108 in the form of tyramine groups. The enzyme horseradish peroxidase (HRP) may be utilized to rapidly catalyze cross-linking of the tyramine groups (e.g., within seconds). Thus, the HRP may be contained within the thermosensitive capsules so that polymerization is not catalyzed until selective and controlled disruption of the capsules in response to applied heat.

[0040] The polymerizable units 102 to which the tyramine groups are attached may be any type of biocompatible, implantable polymer material, including units that can polymerize to form hyaluronic acid, silicone, polyethylene (PE), polyetheretherketones (PEEK), polyethylene terephthalate (PET), polylactide (PLS), polyglycolide (PGA), polytrimethyllenecarbonate (PTMC), poly(p-dioxanone) (PDO), polyurethane (PU), fluoropolymers such as polytetrafluoroethylene (PTFE), hydroxyapatite (HA), hydrogels, and polymethyl methacrylate (PMMA), for example.

[0041] A particularly preferred material for the polymerizable units 102 in soft tissue fillers and/or scaffolding applications is hyaluronic acid. The formation of hyaluronic acid/tyramine conjugates is described in Kurisawa et ak,“Injectable biodegradable hydrogels composed of hyaluronic acid-tyramine conjugates for drug delivery and tissue engineering,” Chem Commun (Camb), 2005(34): p. 4312-4; Lee et ak,“An injectable hyaluronic acid-tyramine hydrogel system for protein delivery,” Journal of Controlled Release, 2009, 134(3): p. 186-93; and Kim et ak,“Injectable hyaluronic acid-tyramine hydrogels for the treatment of rheumatoid arthritis,” Acta Biomater, 2011. 7(2): p. 666-74.

[0042] Other embodiments may utilize other enzymes and/or other polymerizable groups in addition to or as an alternative to the HRP/tyramine system shown in the illustrated biocompatible polymer material.

[0043] As shown in Figures 3A and 3B, the thermosensitive capsules may be formed as liposomes. Liposomes are lipid-based vesicles that self-assemble in aqueous environments. The structure of the liposomes can be disrupted around 40° C, allowing them to release their cross-linking agent contents to thereby catalyze polymerization.

[0044] Figure 3A illustrates a lipid molecule 110 having a hydrophilic head 112 and hydrophobic tail 114. These lipid molecules self-assemble to form liposomes 116 with a lipid bilayer. When formed, the liposomes may encase one or more of the co-mixed cross-linking agents (enzymes 118 in this example). Figure 3B illustrates disruption of a liposome 116 and release of the enzymes 118 upon heating to a temperature of about 40° C (e.g., about 38.5° C to about 43° C). Beneficially, this is a temperature slightly greater than body temperature and is suitable for in vivo applications.

[0045] Liposomes may be synthesized using methods known in the art, such as the hydration methods described in Chandaroy et al.,“Temperature-controlled content release from liposomes encapsulating Pluronic F127,” (in English), Journal of Controlled Release, vol. 76, no. 1-2, pp. 27-37, Sep 11 2001.

[0046] In a preferred embodiment, the thermosensitive liposome is a composite of a lysolipid component and a phospholipid component. The phospholipid component may include one or more of dipalmitoyl phosphatidylcholine (DPPC), distearoyl phosphatidylcholine (DSPC) and hydrogenated soy phosphatidylcholine (HSPC), though other phospholipids may also be included. Preferably, the phospholipid component comprises DPPC. The lysolipid component may include one or both of monopalmitoyl phosphocholine (MPPC) and monostearoyl phosphatidylcholine (MSPC). Preferably, the lysolipid component includes monopalmitoyl phosphocholine (MPPC). Other lysolipids may also be included. In some embodiments, at least a portion of the liposomes may be PEGylated.

In Vivo Cross-Linking via Applied Hyperthermia

[0047] In at least one embodiment, the disclosed controlled cross-linking comprises locally heating the thermosensitive capsules included within the biocompatible polymer material to release the cross-linking agents and initiate polymerization. The energy is focused and steered across the points where the solid is desired. After printing, the remaining un-polymerized materials may be absorbed by the body, leaving behind the in vivo printed structure. Various methods of hyperthermia can be used to provide points of heat to the fluid polymer inside the body, as summarized in Table 1.

Table 1: Heating Methods

Method in/n on-invasive_ resolution

Radiofrequency catheter ablation (RFCA) invasive 1-5 mm

Interstitial microwave heating invasive 0.5-2 cm

Regional focused microwave heating non-mvasive 1-2 cm

Ultra-focused microwave heating non-mvasive < 1 cm [0048] Radiofrequency catheter ablation (RFC A) (typically about 350-500 kHz) is conventionally used to treat cardiac arrhythmias. A coaxial line with the center conductor slightly protruding is used to generate highly localized heating. In at least one embodiment, irrigated RFC A catheters may be used for 3D printing in the body, microfluidically injecting the polymer and simultaneously heating it. RFCA devices conventionally ablate tissue around 80° C. However, adjusting the power allows for a broader range of temperatures, including lower temperatures more suitable for in vivo 3D printing applications (e.g., slightly higher than body temperature, such as about 40° C). Printing with an RFCA probe would be considered minimally invasive surgery.

[0049] Figure 4 illustrates an example distal tip of a probe 120 suitable for use in an in vivo

3D printing application. The probe 120 may form a portion of an RFCA probe, for example. The probe 120 may include one or more lumens extending through its length, and may optionally include one or more perforations 122 to allow fluidic passage of the biocompatible polymer material out of the probe 120 and into the targeted tissue area. The probe 120 is configured to simultaneously heat the extruded polymer material by way of the applied radiofrequency energy.

[0050] In use, the probe 120 can be scanned over the region of the 3D print to form the printed structure. This type of procedure can be controlled either manually, robotically, or magnetically as known in the art of conventional RFCA cardiac arrhyth ia use. RFCA is also compatible with MRI or CT imaging.

[0051] Various different paraments that control the heat pattern including the size/type of probe, frequency of applied energy, amount and type of polymer material extruded from the probe 120, composition of the targeted tissue, blood flow, contact pressure, and time. These parameters can be adjusted according to particular application needs and preferences.

[0052] In another embodiment, microwave hyperthermia may be used to heat and cross- link the biocompatible polymer material. In conventional use for cancer treatment, microwave hyperthermia is used to overheat a cancer tumor, with the aim of minimizing damage to the surrounding healthy tissue. In interstitial microwave hyperthermia, a wire antenna is inserted into a tumor, heating it from the inside out. These methods may be tailored to an in vivo 3D printing application to provide heat for inducing polymerization of the biocompatible polymer material. [0053] A variety of antennas, most derived from a coaxial line, may be utilized for interstitial hyperthermia. Simple monopoles, choked monopoles, segmented monopoles, stepped-impedance monopoles, and sleeved slots may be utilized. Stepped-impedance monopoles can be used to reduce the cold spot that occurs at the tip. Exemplary interstitial applicators that may be utilized in the methods described herein include the BSD-500, provided by Pyrexar Medical Inc. of Salt Lake City. The BSD-500 is an 8-channel, 915 MHz, 480 W system with a variety of applicators including waveguide applicators and microwave interstitial arrays. Another example of a microwave applicator that could be tailored to use for applying the biocompatible polymer material is described in Pisa et ah,“Power density and temperature distributions produced by interstitial arrays of sleeved-slot antennas for hyperthermic cancer therapy,” IEEE Transactions onMicrowave Theory and Techniques , vol. 51, no. 12, pp. 2418- 2426, Dec. 2003.

[0054] The RFCA and/or microwave probes/applicators utilized to deliver the polymer to the targeted anatomical site may include a dielectric sleeve to aid in focusing heating at the tip and to prevent fringing fields along the outer portion of the conductor from producing unwanted tissue heating on the sides of the probe.

[0055] As shown in Figure 5, an array 126 of multiple antennas may also be used to heat a region or bolus 124 of biocompatible polymer material in a focused microwave hyperthermia method. The multiple antennas may be placed around the bolus 124 of biocompatible polymer material. For interstitial applications, the antennas may be placed about 1 to 2 cm apart, for example. The magnitudes and phases of the power fed to the separate antennas 126 can be adjusted to steer and direct heating to various points inside the array boundary 128, including within the targeted bolus 124 to form solidified polymer 130. After forming the desired 3D shape, the remaining fluid of biocompatible polymer material may be safely flushed and/or absorbed by the body.

[0056] As a somewhat related alternative to interstitial arrays, focused regional hyperthermia may utilize one or more antenna arrays located outside the body to focus the fields internally. Focused waves in such an application are typically limited in focal size to about 1/3 to 1/2 wavelength in the material, which is roughly 0.5 to 2 cm in the body. For example, at 915 MHz, deionized water (37° C) has s=0.125 S/m, with a wavelength of 3.806 cm. This is similar to muscle tissue with e_r=55 and s=0.95 S/m. At higher frequencies, the wavelength and focal size are smaller, but so is the depth of penetration, so tradeoffs must be made according to particular application needs. Focus resolution is better near the surface of the body, or in regions that have lower loss (such as the breast or other fatty tissues).

[0057] A wide range of antenna arrays can be used for interstitial or focused regional hyperthermia including linear, planar, hexagonal, circular, cylindrical, spherical, ellipsoidal, interstitial, and annular arrays. Energy can be focused in both the near field and far field using metamaterials, waveguide lens structures, deformable mirrors, prolate spheroidal reflectors, and ultra-wi de-band and time-multiplexed beam formers, for example. Very narrow,“pencil beams” can be used. Several optimization approaches can be applied including an optimal constrained focusing method, iterative techniques, linear optimization, particle swarm, optimized time reversal (OTR) and optimal constrained power focusing technique (OCPF). Arrays of coherent dipoles in a ring pattern or waveguide applicators can also be used.

[0058] Ultra-focused microwave heating may be utilized to target relatively small regions by using highly focused coil designs. For example, MRI receiver 20- and 28-channel imaging coils are capable of focusing on very small anatomical features such as the carotid artery or optic nerve, and thus can be utilized as transmitting antennas to create an ultra-focused and steerable heating pattern. The tradeoff is that such a coil array is ultra-focused only in the specific anatomical region for which it is designed. Improved focus may be obtained with such MRI coils by phasing and combining the returns from each coil independently.

[0059] In at least one embodiment, the system can be used by controlling the phases and magnitudes of each coil independently, creating heat that is more focused than a regional array of other antenna types, albeit steerable over a smaller region of the body. Success of these coils comes from de-coupling from their nearest neighbors. Additionally, similar arrays of crossed dipoles, which may have better de-coupling, can be used.

[0060] Various contrast-enhancing and/or heat-enhancing materials may also be utilized such as magnetic materials, nanoparticles (e.g., gold), nanotubes, and microbubbles. These materials may be added to the biocompatible polymer material in order to make it better absorb applied energy and heat more quickly than nearby tissue. Iron oxide magnetic nanoparticles are particularly preferred. Iron oxide nanoparticles have a facile and mild synthesis in an aqueous solution, which can be easily incorporated into an aqueous biocompatible polymer material prior to in vivo polymerization. These nanoparticles are biocompatible, and their degradation mechanism in the body is well-understood. Additionally, or alternatively, implanted low-frequency electrodes and thermal seed heating can also be used. [0061] Some embodiments may utilize ultrasound to provide focused in vivo heating. However, ultrasound is less preferred than other electromagnetic energy methods because typical biocompatible polymer materials are unable to attenuate sound waves to a degree that allows sufficient heating.

[0062] One will appreciate that the human body is a highly heterogeneous structure, with electrical properties varying from tissue to tissue. Energy delivered from antennas in or near the body depends on the anatomy and size/shape/tilt/location of the antenna, the size and shape of the body (e.g., adults vs. children), and variation in tissue properties (e.g., from person to person or just because of uncertainty in the measurements). In addition, the heat created from the delivered energy may depend on the blood flow and thermal properties of the surrounding tissues. Temperature and/or delivered energy levels are therefore preferably monitored in real time to account for particular application needs and differences.

Conductive Biocompatible Nanocomposites

[0063] Certain biocompatible polymer materials described herein may also be formulated to be conductive. For example, conductive nanoparticles may be embedded in a matrix of biocompatible polymer material to provide a material that has high electrical conductivity but low mechanical stiffness. The volume ratio of conductive nanoparticles in the polymer matrix may be adjusted to balance and control electrical conductivity and mechanical properties of the nanocomposite. Conductive biocompatible polymer materials formed in this manner may then be utilized for in vivo 3D printing and/or heat-activated cross-linking as described above in relation to other embodiments.

[0064] Conductive biocompatible nanocomposites provide several benefits. For example, many conductive implants are made of metal to provide sufficient electrical conductivity. However, metal implants may have limitations related to biocompatibility and the ability to match the mechanical properties of the surrounding tissues. For example, the Young’ s modulus of soft tissues is usually on the order of kPa, but the Young’s modulus of gold is about 78 GPa. The mismatch of the mechanical properties between the tissue and metal can cause trauma and/or induce inflammation. The conductive biocompatible nanocomposites described herein effectively tempers this mismatch to provide mechanically functional yet sufficiently conductive biomaterials.

[0065] In one embodiment, for example, the electrical conductivity of the nanocomposite is on the order of about 10⁶ to 10⁷ S/m, but with a Young’s modulus that is an order of magnitude less than gold. Such an embodiment may include gold nanoparticles in a cationic polymer such as cationic polyurethane, though any of the other polymer systems described herein may also be utilized, including those described above formulated to be enzymatically cross-linkable.

[0066] The nanoparticles may be formed using chemical synthesis methods such as reduction methods (e.g., the citrate reduction method). Nanoparticles may additionally or alternatively be formulated using a laser ablation process or other non-chemical manufacturing process. Preferred embodiments utilize gold nanoparticles. However, other embodiments may additionally or alternatively utilize one or more other conductive metal nanoparticles, such as those comprising platinum, silver, aluminum, tin, tungsten, chromium, titanium, magnetite, charcoal, hematite, cinnabar, limonite, corundum, rutile, zincite, ferric ferrocyanide, cobaltous aluminate, or combinations thereof.

[0067] The nanoparticle density of the conductive nanocomposite may be adjusted to provide desired balance between mechanical properties of the material and the electrical conductivity of the material. In some embodiments, the nanoparticles may be provided at about 0.1 to about 35 g per cm³ of the nanocomposite, or about 1 to 25 g per cm³ of the nanocomposite, or about 5 to 20 g per cm³ of the nanocomposite, though these ranges may vary according to particular application needs.

3D Printable Antennas for Implantable Medical Devices

[0068] Conductive nanocomposite materials as described above may be utilized to form effective antenna implants for use with IMDs. Such antennas may be solid, segmented, meshed, or formed in other appropriate structures. IMDs often need to send and/or receive electromagnetic fields in order to enable wireless telemetry and/or WPT. An IMD may operate at various operational frequencies such as 402 MHz, 433 MHz, 868 MHz, 915 MHz, 1.4 GHz, 2.45 GHz, 6 GHz, or other frequencies within a range having endpoints defined by any two of the foregoing values. WPT methods may include near-field resonant inductive coupling (NRIC), near-field capacitive coupling (NCC), midfield WPT, and far-field electromagnetic coupling (FEC).

[0069] Modem implantable antennas are limited by the size of the obj ect they can be placed on. Pacemakers and other such IMDs are typically on the order of a few square centimeters, but continue to decrease in size as battery and communication technology becomes progressively more sophisticated. In some applications, only the feed system for the antenna is placed on the implant, and is passively coupled a short distance through the body to the surface (i.e., skin), where a larger antenna can pick up the signal and re-radiate it. A preferred location for such antennae is in the fat layer beneath the skin (to avoid shorting effects of more conductive muscle tissue), though some may be placed in or extend into deeper tissue layers (e.g., muscle).

[0070] Other applications where a biocompatible conductive material is desired may utilize the described biocompatible and conductive nanocomposites. The conductive nanocomposites may be formed into wires, plates, and 3D materials, for example, to provide various conductive functions. Some embodiments may directly connect the implantable device with the surface and/or a surface antenna via a subsurface wire formed from a conductive nanocomposite. Any other electrical device (sensor, capacitor, inductor, etc.) could also be 3D printed within the body with a biocompatible conductive nanocomposite material as described herein.

[0071] For instance, a wire-like antenna may be implanted using a catheter or needle and retracting it as polymer is extruded to produce an internal wire. In some embodiments, heat may also be applied simultaneously or subsequently to induce polymerization, as described above. Other antennas or electrodes may be formed with a plate shape or with another 3D structural shape.

[0072] Figure 6A is a cross-sectional view of tissue showing an embodiment of a 3D printable, dipole nanocomposite antenna that includes two wires 132 and is configured to guide fields (e.g., generated from a far-field source) from the surface 16 (i.e., skin) to deeper regions of the body. An IMD may be located at or near the focused zone 134. The small gap at the focused zone 134 between the tips of the lines provides constructive coupling and augments the focused field. The focused fields may be used for recharging the battery of an IMD as a WPT function or they may enable an IMD to communicate with an external source as a telemetry function.

[0073] Figure 6B illustrates a similar embodiment including multiple pairs of wires 132 of different length values. The use of multiple pairs (e.g., within the same plane) can increase the specific absorption rate (SAR). As compared to the case where no such antenna lines are used, the power of the focused field may be up to about 8 times to about 16 times larger where one or two lines are utilized, respectively. In the case where multiple lines are utilized, as shown in Figure 6B, the power of the focused field may be up to about 20 times larger than a no wire baseline. [0074] The wires 132 may have various lengths suitable for different application needs. Some wires 132 may have lengths of about 20 mm to about 70 mm, or about 30 mm to about 60 mm, for example, though values may be customized according to IMD depth, type of tissue in which the wires pass (e.g., muscle vs. fat), and the like. Where one or more wire pairs are utilized, the gap may range in size from about 1 mm to about 8 mm, or more typically about 1.5 mm to about 4 mm. Generally, smaller gap sizes of about 2 mm or less are preferred in order to provide more effective focusing. However, where other design constraints necessitate larger gaps such larger gaps may still be utilized to provide effective focusing. The depth at which the gap is positioned will also vary according to particular IMD type and application specifics, but may generally be about 10 mm to about 60 mm from the skin surface 16.

[0075] Figure 6C illustrates another embodiment where two pairs of wires 132 (first pair l32a and second pair l32b) are arranged in a crossed-dipole configuration with each pair positioned orthogonal to the other. Figure 6C is shown in plain view looking down at the skin surface 16. This type of configuration may be beneficially utilized in for transmitting and/or receiving circularly polarized waves.

[0076] Although the wired ends that form the coupling gap (i.e., the ends that are positioned deepest in the tissue) are shown here as being substantially aligned, it will be recognized that there may be some level of misalignment in real world applications due to imperfect implantation/printing technique, anatomical quirks, or other application imperfections. Beneficially, the wire antenna embodiments have been found to be tolerant of some degree of misalignment, being able to provide effective focusing even where such misalignment exists. For example, if the depth of the distal ends of the wires (i.e., the ends located deepest in the tissue) in a wire pair are offset by a range of about 2 mm to about 8 mm, the corresponding deterioration of the SAR ranges from about 1% to about 30%. While not ideal, the ability to maintain good focusing capabilities even with depth misalignment illustrates the versatility and robustness of the design.

[0077] Although the examples shown in Figures 6A through 6C angle the wires 132 such that the distal ends are deeper within the tissue than the proximal ends, other embodiments may utilize antennae that are not angled. For example, some antennae may position wires, wire pairs, or wire arrays with the wires oriented substantially parallel to the skin surface.

[0078] Figures 7A through 8B illustrate other embodiments of in vivo 3D printable antennae formed as rings. The rings may be shaped as circles (Figures 7A and 7B) or ellipses (Figures 8A and 8B). Preferably, the ring antennae have a 3D arrangement, with concentric rings being set at different depths. For example, as shown, each successive inner ring is positioned at a greater depth to form a“funnel” for focusing applied electromagnetic waves at the focused zone 234/334. In other embodiments, however, the antennae have a 2D arrangement with concentric rings positioned along the same plane (i.e., at the same depth).

[0079] Figure 7A shows a plan view of the circular shaped rings 232, from a position looking down at the skin surface 16, and Figure 7B shows the same rings 232 in a cross- sectional view. Figure 8 A shows a plan view of the ellipse shaped rings 332, from a position looking down at the skin surface 16, and Figure 8B shows the same rings 332 in a cross- sectional view.

[0080] Ring shaped antennae as described herein have been found to increase SAR at the desired focused zone 234/334 by a factor of about 2 to 4 as compared to a baseline not having any antennae. Circular shaped rings are more preferred for circular polarization applications while ellipse shaped rings are more preferred for linear polarization applications. An ellipse shaped ring arrangement was found to focus and enhance fields inside muscle tissue with a much smaller area than a circular shaped ring arrangement, though with somewhat higher increases in SAR.

[0081] The number of concentric rings utilized may be varied according to particular design needs, IMD requirements, and cost considerations. A single ring may be utilized or an array of multiple concentric rings may be utilized having up to about 8 concentric rings, or more commonly about 2 to 5 concentric rings, for example.

[0082] Ring size may likewise be varied according to particular design needs, IMD requirements, and cost considerations. Ring diameter may be within a range of about 20 mm to about 200 mm, or more commonly about 50 mm to 150 mm. Where multiple concentric rings are utilized together in an array, the largest, outermost ring may have a diameter that is about 30% to about 70% greater than the smallest, innermost ring, or about 40% to about 50% greater than the smallest, innermost ring. For ellipse shaped rings, similar values apply. For example, the major axis of the ellipse may have a length of about 20 mm to about 200 mm, or more commonly about 25 mm to 100 mm, and the minor axis may have a length that is about 15% to about 80% of the major axis, or more commonly about 20% to about 50% of the major axis. [0083] As described above, the wire antennas or other conductive electronic components may be formed by injecting and simultaneously or subsequently heating the nanocomposite to induce polymerization. Heating may be accomplished using the methods described above. For wire antennas in particular, a preferred method includes the steps of injecting the conductive nanocomposite through a catheter, needle, and/or RFCA probe while also exciting it with radiofrequency energy. As the polymer is injected and heated, the wire can be backed out, leaving behind a polymerized, solid antenna structure. After this 3D printing process, remaining fluids, if any, will be naturally washed or absorbed by the body.

[0084] The thermally polymerizable materials utilized in the biocompatible nanocomposites described herein may include any of the enzymatic-based cross-linking formulations described above, or any other biocompatible polymer capable of retaining sufficient fluid properties at body temperature but polymerizing at temperatures slightly above body temperature (e.g., 38.5° C to about 43° C). Cellulose ethers are one class of polymers that may be utilized, for example.

[0085] As used herein, the terms“about,”“approximately,”“substantially,” and the like may be taken to refer to an amount that deviates by no more than about 10%, or by no more than about 5%, or by no more than about 2.5%, or by no more than about 1% of the stated value or condition.

EXAMPLES

Example 1

[0086] Gold nanoparticles were synthesized using a citrate reduction method. 90 mg of HAuCl4 is dissolved in 500 ml of water. The solution was heated on a hot plate until boiling. Then 25ml of 0.1% sodium citrate aqueous solution was added to the gold salt solution. The mixture was stirred and reboiled on the hot plate. After 20 minutes, the solution became a red color, which indicated the formation of Au NPs.

Example 2

[0087] lml of 1.0 vol.% aqueous cationic polyurethane was slowly added to a liter of gold nanoparticle dispersion as prepared in Example 1. The mixture was stirred for 15 min followed by filtration. Filter papers were of 0.8-pm pore size with 47mm diameter. The resultant gold- colored nanocomposite was readily peeled off the filter paper.

Example 3 [0088] Dipole wire Antennae formed from the nanocomposite of Example 2 were placed in various locations in pork tissue. Placing an antenna in or on top of the fat layer was seen to reduce shorting effects significantly, whereas similar antennae placed on or in muscle tissue showed significant detuning due to short circuiting.

Example 4

[0089] A dipole antenna formed from the nanocomposite of Example 2 was compared against several other different materials in a porcine tissue model. The results are shown in Table 2. Three different gold nanocomposite materials were formed by mixing various concentrations of polyurethane and gold.

Table 2: Material Conductivities

[0090] The antenna made from copper tape served as a benchmark. The antenna made from the gold nanocomposite was biocompatible, highly conductive, and flexible. The other comparison antennas were made from conductive inks. Although they are not biocompatible themselves, they provided proxies for other materials that could be utilized for subdermal injection. The NovaCentrix material is a conductive ink that can be inkjet printed with small (250 pm) feature size onto a sheet of thin PET plastic. PELCO product #16062 is a conductive paint from Ted Pella Co. that is easily spread. The thickness of the copper tape and PELCO is significantly greater than the skin depth (assumed as 30.4 mm in this example), the gold nanoparticle polymer is close to the skin depth, and the NovaCentrix material is thinner than the skin depth.

[0091] Additional testing showed that antennas insulated in by the fat layer have better Su parameter and higher resonant frequency than those in direct contact with muscle.

Example 5 [0092] Field focusing effects of passive conductive wires formed from a biocompatible nanocomposite were modeled. The tissue section of the human body in which the wires were placed was modeled as a 3D stratified medium with the layer parameters given in Table 3.

Table 3: Material Properties at /= 433 MHz

[0093] The model positioned the desired IMD location at a depth of 30 mm. The wires/lines were modeled as flat tipped cylinders with diameter of 0.2 mm, length of 40 mm, and gap between distal/deep wire ends of 2 mm. The cross-sectioned surface area of the layered body model was 180 x 180 mm². Input power was set at 0.82 W.

[0094] Compared to baseline (no wires implanted), the model showed that a single wire increased SAR by a factor of about 8 at the desired IMD location, while a wire pair (such as shown in Figure 6B) increased SAR by a factor of about 16 at the desired IMD location. Increasing the number of line pairs to a six-line design (such as shown in Figure 6C) increased SAR by a factor of about 20 at the desired IMD location. Results are summarized in Table 4.

Table 4: Comparison of SAR for Focused Line Designs (Pin = 0.82 W)

Example 6

[0095] Field focusing effects of conductive ring arrays formed from a biocompatible nanocomposite were modeled. The tissue section of the human body in which the wires were placed was modeled as a 3D stratified medium with the same layer parameters for skin, fat, and muscle given in Table 3 above. Results are summarized in Table 5. Table 5: Comparison of SAR for Ring Arrays (Pin = 0.82 W)

[0096] The modeled arrays included three rings each. Optimal depths of the rings of the circular shaped ring array were determined to be 10 mm, 18.88 mm, and 19.21 mm, with respective radii of 57.91 mm, 39.68 mm, and 31.63 mm. Optimal depths of the rings of the ellipse shaped ring array were determined to be 10 mm, 17.15 mm, and 17.21 mm, with respective minor/major axes of 13.7/48.74 mm, 10.32 /32.07 mm, and 5/23.43 mm.

Claims

1. A method of printing an implant in vivo within a subject, the method comprising: applying a biocompatible polymer material to a targeted anatomical region, the biocompatible polymer material being formulated so as to remain unpolymerized at body temperature and to be polymerized when exposed to sufficient heat; and

applying heat to at least a portion of the biocompatible polymer material to thereby cause at least a portion of the biocompatible polymer material to polymerize and form a solid structure at the targeted anatomical region.

2. The method of claim 1, wherein the targeted anatomical region is soft tissue.

3. The method of claim 1 or claim 2, wherein the biocompatible polymer material is injected into the targeted anatomical region.

4. The method of any one of claims 1 through 3, wherein the formed solid structure is a scaffold, support structure, filler, or soft tissue replacement.

5. The method of any one of claims 1 through 4, wherein the soft tissue replacement replaces one or more of cartilage, ligaments, and tendons.

6. The method of any one of claims 1 through 5, wherein the targeted anatomical region includes the face as part of a facial reconstruction surgery.

7. The method of any one of claims 1 through 6, wherein the formed solid structure is an antenna implant.

8. The method of claim 7, wherein the antenna implant is configured to work in conjunction with an IMD to enable and/or enhance wireless telemetry and/or WPT to the IMD.

9. The method of claim 8, wherein the IMD is a retinal implant, cochlear implant, glucose monitor, lactate concentration monitor, strain sensor, temperature monitor, or pacemaker.

10. The method of any one of claims 1 through 9, wherein the biocompatible polymer material comprises polymerizable units and a cross-linking agent encapsulated in thermosensitive capsules.

11. The method of claim 10, wherein the thermosensitive capsules are formulated so as to degrade at about 40° C.

12. The method of claim 10 or claim 11, wherein the polymerizable units are enzymatically cross-linkable and wherein the cross-linking agents are enzymes.

13. The method of claim 12, wherein the polymerizable units include tyramine groups, and wherein the enzymes are horseradish peroxidase enzymes.

14. The method of any one of claims 10 through 13, wherein the polymerizable units polymerize to form hyaluronic acid.

15. The method of any one of claims 1 through 14, wherein the thermosensitive capsules are liposomes.

16. The method of claim 15, wherein the liposomes are lysolipid/phospholipid composites.

17. The method of any one of claims 1 through 16, wherein heat is applied to the biocompatible polymer material using radiofrequency and/or microwaves.

18. The method of claim 17, wherein heat is applied to the biocompatible polymer material using one or more of radiofrequency catheter ablation, interstitial microwave heating, regional focused microwave heating, and ultra-focused microwave heating.

19. The method of any one of claims 1 through 18, wherein a probe utilized to apply the biocompatible polymer material is also configured to simultaneously heat the biocompatible polymer material as it is extruded from the probe.

20. The method of any one of claims 17 through 19, wherein an array of antennas is utilized to direct heat to targeted regions within the biocompatible polymer material.

21. The method of any one of claims 1 through 20, wherein the biocompatible polymer material is a conductive nanocomposite including conductive nanoparticles.

22. The method of claim 21, wherein the nanoparticles comprise gold nanoparticles.

23. The method of claim 21 or claim 22, wherein the conductive nanocomposite has a conductivity of about 10⁵ to 10⁷ S/m.

24. The method of any one of claims 21 through 23, wherein the conductive nanoparticles are included at about 0.1 to about 35 g per cm³ of the nanocomposite, or about 1 to 25 g per cm³ of the nanocomposite, or about 5 to 20 g per cm³ of the nanocomposite.

25. The method of any one of claims 21 through 24, wherein the formed solid structure is an antenna implant.

26. The method of claim 25, wherein the antenna implant is a dipole antenna of two paired lines.

27. The method of claim 26, wherein each line of the antenna implant is angled such that the distal ends approach one another and are disposed deeper in tissue than the proximal ends.

28. The method of claim 26 or claim 27, further comprising a second pair of lines positioned orthogonal to the first pair of lines.

29. The method of claim 28, wherein the second pair of lines are each angled such that the distal ends approach one another and are disposed deeper in tissue than the proximal ends.

30. The method of claim 25, wherein the antenna implant includes an array of ring-shaped antennas.

31. The method of claim 30, wherein the ring-shaped antennas include circle-shaped antennas.

32. The method of claim 30 or claim 31, wherein the ring-shaped antennas include ellipse- shaped antennas.

33. The method of any one of claims 30 through 32, wherein the ring-shaped antennas have varying sizes and are disposed in a concentric configuration.

34. The method of claim 33, wherein the concentric configuration positions successively smaller ring-shaped antennas deeper in tissue.

35. A medical implant formed using a method as in any one of claims 1 through 34.

36. A biocompatible polymerizable material formulated to polymerize upon exposure to sufficient heat and being capable of polymerization in vivo , the material comprising:

a polymerizable component that includes a plurality of polymerizable units;

a plurality of thermosensitive capsules; and

a cross-linking agent encapsulated within the thermosensitive capsules, the cross- linking agent being configured to induce cross-linking of the polymerizable units,

wherein the thermosensitive capsules are configured to hold the cross-linking agent at body temperature and to release the cross-linking agent upon heating to a temperature above body temperature.

37. The material of claim 36, wherein the thermosensitive capsules are formulated so as to release the cross-linking agent at about 40° C.

38. The material of claim 36 or claim 37, wherein the polymerizable units are enzymatically cross-linkable and wherein the cross-linking agents are enzymes.

39. The material of claim 38, wherein the polymerizable units include tyramine groups, and wherein the enzymes are horseradish peroxidase enzymes.

40. The material of any one of claims 36 through 39, wherein the polymerizable units polymerize to form hyaluronic acid.

41. The material of any one of claims 36 through 40, wherein the thermosensitive capsules are liposomes.

42. The material of claim 41, wherein the liposomes are lysolipid/phospholipid composites.

43. A biocompatible, conductive nanocomposite material, comprising:

a polymer matrix formed from the material as in any one of claims 36 through 42; and a plurality of conductive nanoparticles dispersed within the polymer matrix.

44. The nanocomposite of claim 43, wherein the conductive nanoparticles comprise gold nanoparticles.

45. The nanocomposite of claim 43 or claim 44, wherein the conductive nanoparticles are included at about 0.1 to about 35 g per cm³ of the nanocomposite, or about 1 to 25 g per cm³ of the nanocomposite, or about 5 to 20 g per cm³ of the nanocomposite.

46. The nanocomposite of any one of claims 43 through 45, wherein the conductive nanocomposite has a conductivity of about 10³ to 10⁷ S/m, or about 10⁴ to 10⁶ S/m.

47. A biocompatible, conductive nanocomposite material, comprising:

a polymer matrix; and

a plurality of conductive nanoparticles dispersed within the polymer matrix.

48. The nanocomposite of claim 47, wherein the conductive nanoparticles comprise gold nanoparticles.

49. The nanocomposite of claim 47 or claim 48, wherein the conductive nanoparticles are included at about 0.1 to about 35 g per cm³ of the nanocomposite, or about 1 to 25 g per cm³ of the nanocomposite, or about 5 to 20 g per cm³ of the nanocomposite.

50. The nanocomposite of any one of claims 47 through 49, wherein the conductive nanocomposite has a conductivity of about 10³ to 10⁷ S/m, or about 10⁴ to 10⁶ S/m.