Authors: Luís Martins, Ana Isabel Barbosa, Vítor Manuel Correlo, Mrinal Bhattacharya, Rui Luís Reis
Categories: Review Article
Source: Bioactive Materials
Authors: Luís Martins, Ana Isabel Barbosa, Vítor Manuel Correlo, Mrinal Bhattacharya, Rui Luís Reis
Piezoelectricity refers to the phenomenon in which certain materials convert mechanical energy into electrical energy and vice versa. It occurs in natural and synthetic materials, underscoring its broad importance within biological environments. This review will discuss the basic principles of piezoelectricity, with a focus on its presence in synthetic materials, including polymers like polyvinylidene fluoride and other polyesters, along with composites of these polymers. The review will also highlight the natural piezoelectric responses observed in human tissues, including bone, skin, dental tissues, and connective tissues and relate these effects to their non-centrosymmetric molecular structure. Traditional tissue engineering materials focus primarily on biochemical, mechanical signals without sustaining the complexity of a natural microenvironment. Piezoelectric materials may offer a new approach to tissue engineering, providing electrical signals capable of directing cellular behavior. These mechanical generated signals are able to create a dynamic and self-powered method for enhancing cellular communication, survival, and differentiation, particularly applicable to regenerative strategies in bone and neural tissue. The review will also consider the most recent discoveries around the use of piezoelectric materials in the scaffolding systems supporting the growth of bone and nerve tissues. and their use in repairing skin and skeletal muscle, showing potential utility in even broader regenerative applications, while also importantly highlighting the adaptable nature of piezoelectric materials.
The piezoelectric effect, first described in 1880 by Pierre and Jacques Curie in crystalline quartz, describes the ability of certain materials to generate an electric charge under the application of mechanical stress. The term piezoelectricity is derived from the Greek words piezein (to press) and elektron (amber, known to be a source of electric charge). Notably, the effect is reversible; a superimposed external electric field can draw out a mechanical strain of the material, a phenomenon called the converse piezoelectric effect [1] (Fig. 1a). Direct and converse piezoelectric effects arise due to the interplay between electric and mechanical conditions of materials lacking symmetry [2]. Upon deformation in these materials, positive and negative centers of charges get displaced, resulting in the creation of dipoles and polarization of the system, accompanied by the development of an electric field. Such an electromechanical coupling enables the conversion of mechanical and electric energy and vice-versa, and thus, piezoelectricity becomes one of the most versatile material attributes of the modern age of science and engineering.Fig. 1a) Schematic of direct piezoelectric effect (left) and converse piezoelectric effect (right). b) Coordinate systems and direction of polarization/electric field of piezoelectric materials c) Longitudinal expansion/Contraction mode. d) Transverse expansion/Contraction mode. e) Parallel shear mode. f) Bending mode. Reproduced with permission from Ref. [13]. g) Diagrams of the α, β and γ polymorphic phases of PVDF. Reproduced with permission from Ref. [69].Fig. 1
Previously reserved only for crystals and ceramics like quartz and lead zirconate titanate (PZT), the area of piezoelectricity was recently revolutionized through findings of piezoelectricity in polymers (e.g., PVDF and copolymers) and in natural tissue materials like bone, tendon, and collagen fibrils. While soft and biocompatible, unlike hard ceramics, polymers are mechanically soft, light in weight, and chemically tailorable. Most importantly, soft and biocompatible, this new area of materials expands the design space of piezoelectric materials, moving toward technologies ranging from wearable electronics and self-powered health monitors, through implantable biomedical devices, back again toward new types of energy-harvesting fabrics and clothing. Of further significance, nanoscale processing and molecular tailoring now enable scientists and engineers, at the molecular level, to control chain alignment, crystal structure, and dipole orientation, and thus, at the macro level, substantially increase the piezoelectric response of soft materials.
Despite these advances, sweeping gaps in scientific and technological knowledge remain. Very little is satisfactorily accounted for in terms of molecular orientation, nanometer-scale structure, and long-range order control of the piezoelectric response of polymers and bio-based materials. Further, device-level integration of these materials into robust, scalable, and multi-functional devices challenges vigorously and awaits intensive study. Redressing these shortcomings holds a singular promise of not only enhancing our fundamental understanding of electromechanical coupling in soft materials but also encouraging innovation in energy scavenging, sensing, and biomedical devices.
While recent reviews have advanced the understanding of piezoelectric biomaterials, they often remain segmented by tissue type [3] or material system or even specific mechanisms [4], for instance, focusing on bone regeneration [5] or neural repair [6]. Other studies emphasize material fabrication strategies [7] and application-specific outcomes [8] rather than exploring the underlying physical principles and cross-tissue approaches.
In contrast, the present review provides a theory-driven and integrative approach by synthesizing the piezoelectric phenomenon from its quasi-electrostatic foundations and molecular origins in non-centrosymmetric structures to its manifestation across multiple biological tissues, including bone, skin, dental, and connective tissues. It further systematically evaluates synthetic polymers such as PVDF and their composites for broad tissue engineering applications, encompassing bone, cartilage, neural, muscle, and skin repair. By combining theoretical understanding with recent experimental insights and design principles, this work establishes a unified framework and identifies new pathways for scalable, multifunctional scaffolds in regenerative medicine, thereby extending beyond the scope of existing reviews.
This current review is stimulated by the remarkable transformative capability of piezoelectric materials based on polymers. It systematically compares and contrasts various classes of materials, demonstrating how processability, biocompatibility, and tailored mechanical flexibility make polymers uniquely suited for diverse tissue engineering applications. It would also explore the use of polymer-ceramic composites, where polymers are used as the continuous phase to provide flexibility while incorporating ceramic nanoparticles to boost the overall piezoelectric effect.
By bringing together understanding at the molecular level with new processing and modeling approaches, we aim in this White Paper to define design principles that enable the systematic creation of next-generation technologies enabled by piezoelectricity. Prospective deliverables carry the promise of revolutionizing the role of piezoelectricity in society tomorrow, transitioning it from a property of stiff crystalline solids, through a multi-functional, shape-variable capability, built into wearable, conformable, and biologically interfaced technologies.
The piezoelectric theory is derived using quasi electrostatic approximations (assuming that the magnetic effects are negligible compared to electrical effects) thus avoiding the use of the complete set of electromagnetic equations. The Polarization density P (Cm^−2^) upon the application of an electric field E (Vm^−1^) on a dielectric is given (1)P=χeEwhere χ is the dielectric susceptibility (C2Nm2) is a second ranked tensor. Dielectrics are insulators that are polarized when subjected to electric field. The imposition of an electric field on a dielectric causes the charges to separate inducing a electric dipole moment. For linear dielectrics χ is a constant. The electric displacement field D is given by(2)D=ε0E+Phere D is a first order tensor, and ε0 is the dielectric permittivity of a vacuum. It can be shown that [9] D = εE where ε = ε0εr is the dielectric permittivity of the material and εr = 1+χ is the relative permittivity of the material.
Thus, the linear piezoelectric theory involves coupling the linear elasticity equation with the electrostatic equation by means of the piezoelectric constants. The relationship between stress and strain for an elastic material is given by Hooke's law(3)τij=cijklSkl
The conservation of angular momentum necessitates that the tress tensor be symmetric [10]. Hence, τij=τji and Skl=Slk. The symmetry also necessitates that cijkl=cjikl and the number of independent elements of the stiffness tensor is reduced from 81 to 21 [11,12]. The symmetry of the stress and strain tensors necessitates the symmetry of piezoelectric constants, i.e. dikj=dkij. This reduces the number of independent piezoelectric constants from 27 to 18. For piezoelectric material, the relationship between the stress τij and the electric field displacement is given by(4)Di=dijkτijwhere d the piezoelectric constant is charge generated per applied force or deflection per applied voltage. The piezoelectric constant can have normal and shear components. In addition, dimensions of piezoelectric materials are affected upon the application of an electric field. This is given by(5)Sik=dijkTEk=dkijEkwhere T denotes the transpose of the matrix. It should be noted that D and E are vectors, S and τ are second order tensor, and the piezoelectric constant is a third order tensor. The piezoelectric constant is an indication of how well a material can convert a mechanical force into an electric output. This constant can assume a positive or a negative value. The sign associated with the displacement D and strain Skl is dependent on the direction of the respective mechanical and electrical field [9,13]. Since piezoelectric materials are anisotropic each constant has two subscripts; one that relate to the direction of the stress and the other to the direction of strain (Fig. 1b). The normal direction is represented by subscripts 1, 2, and 3, while the shear direction is represented by subscripts 4, 5, and 6.
The piezoelectric constitutive equations as outlined in the IEEE standards [14] in the stress-charge form is(6)τij=cijklESkl−ekijEk(7)Di=eiklSkl+εikSEkHere τij is a 6 x 1 column vector, cijklE is a 6 x 6 matrix, Skl is a 6 x 1 column vector, ekij is 6 x3 matrix, and E is 3 x 1 column vector. Similarly, D is 3 x 1 column vector, ε is a 3 x 3 matrix.
The strain-charge form of the constitutive equations are(8)Sij=sijklEτkl+dkijEk(9)Di=diklτkl+εikτEk
It should be noted that the elastic constant cijklE has 21 independent constants, the piezoelectric constant ekij has 18 independent constants, and the permittivity constant εij has six independent constants.
There are alternate forms of the constitutive equations listed in the IEEE standards. For converse piezoelectric effect the set of constitutive equations are(10)Sij=sijklDτkl+gkijDk(11)Ei=−giklτkl+βikτDk
and,(12)τij=cijklDSkl−hkijDk(13)Ei=−hiklSkl+βikSDk
A compressed matrix notation is introduced in place of tensor notation so that the elastic compliance or piezoelectric constant is replaced by matrix notation. The pairs of repeating indices ii = 11, 22, 33 is replaced p or q = 1, 2, 3 and the shear and strain components represented by the mixed indices such as ij or kl where ij = 12 or 21, 23 or 32, 31 or 13 is replaced by p or q as 3, 4, 5 [14]. The reduced constitutive equations related to the stress charge form (eq. (6) and (7)) becomes(14)τp=cpqESq−ekpEk(15)Di=eiqSq+εikSEk
The strain charge form reduces (16)Sj=sjlτl+djiEi(17)Di=dijτj+εikEk
The four piezoelectric coefficients dij, eij, gij, and hij are defined (18)dij=(∂Di∂τj)E=(∂Sj∂Ei)τ(19)eij=(∂Di∂Sj)E=−(∂τj∂Ei)S(20)gij=−(∂Ei∂τj)D=(∂Sj∂Di)τ(21)hij=−(∂Ei∂Sj)D=−(∂τj∂Di)S
The slopes in equations (18), (19), (20), (21)) are evaluated at the constant value of the superscripts (D, E, S, and τ). The first set of terms corresponds to the direct piezoelectric effect while the second set of terms corresponds to the converse piezoelectric effect. In the suffix ij, i represents the direction of the electric field and j represents the direction of mechanical action.
The above sets are phenomenological equations that were originally derived for ceramics which possesses a regular crystal structure. They are thus treated as linear phenomena i.e. the material properties are invariant of both applied stress and electric field. However, polymers have different structures in that they are often semi-crystalline, with crystals dispersed in an amorphous phase. The validity of these equations for polymers have not been rigorously tested. Indeed, there has been evidence that piezoelectric stress constant are a nonlinear function of the applied stress or strain for poly(vinylidene fluoride) [15].
Polymers can be classified as either polar or non-polar. Polar molecules have uneven electron distribution resulting in dipole moments and exhibit high dielectric constant. In contrast, non-polar molecules have evenly distributed molecules where the dipole moments because of their symmetry cancel out. Interestingly, biological tissues such as wool and hair under stress were shown to display piezoelectric effects [16,17] due to the presence of chiral atomic groups [18]. Later, studies indicated both natural, synthetic, and optically polymers displayed piezoelectric properties [19].
Most polymers display piezoelectric properties [2,20] on account of its structure and orientation. These are summarized in Table 1. The list includes fluoropolymers [[21], [22], [23], [24]], polyurea's [[25], [26], [27]], polyamides such as odd numbered nylons [[28], [29], [30], [31], [32], [33], [34]], elastomers [35], poly-L-lactic acid [36,37], polyhydroxy butyrate and their copolymers [38,39]. Natural polymers such as collagen [[40], [41], [42]], DNA [18,43], bone [[44], [45], [46], [47], [48], [49]], silk [50,51], and others [43,[52], [53], [54]] also exhibit piezoelectricity. These connective tissues often consist of fibrous molecules that has crystalline structure. There is similarity between these tissues and liquid crystals [55]. A recent study [56] gives a chronological sequence of research in ferroelectricity. Notably, ferroelectric materials are a subset of piezoelectric materials, meaning, all ferroelectric materials are piezoelectric, while the reverse is not true.Table 1Effect on orientation on piezoelectric properties of polymers.Table 1MaterialOrientationMethodStress Frequencyd14 (pC/N)ReferencePoly-α-isobutyl-L-aspartateUnorientedStatic0[277]4 Hz0Poly-α-isobutyl-L-aspartateElectric fieldStatic13[277]4 Hz11.2Poly(γ-methyl-L-glutamate)Stretching10 Hz2[18]Poly(γ-methyl-L-glutamate)Magnetic Field10 Hz4[18]Poly(γ-methyl-L-glutamate)Rolling10 Hz2[18]SilkStretching[536]λ = 10.01λ = 2.71.5CollagenOriented2–3[18]
PVDF (−CH2−CF2)n is the most studied polymer in terms of its piezoelectric characteristics. It is a semi crystalline and inherently polar polymer. Piezoelectricity in uniaxially drawn and poled PVDF films was first reported by Kawai [21]. Since then, extensive research has been conducted to understand the fundamentals behind the piezoelectric, ferroelectric, and pyroelectric behavior of PVDF [[57], [58], [59], [60]]. Initially, there was some debate about the origin of piezoelectricity. It was suggested that PVDF was a ferroelectric polymer as evidenced by the dipole reorientation during poling [22,61,62]. Ferroelectricity in PVDF polymers were verified from dielectric hysteresis (D-E) loop measurement [[62], [63], [64]]. The electrical characteristics are due to the positively charged hydrogen and the negatively charged fluorine. Upon cooling of the melt, it crystallizes into spherulite domains. It has a low glass transition temperature (Tg) of approximately −35^0^C, while its melting temperature (Tm) is around to 175^0^C. The volume fraction of crystal ranges from 50 % to 70 % and depends upon the thermal history. Amongst polymers, PVDF has one of the highest piezoelectric responses [65].
There are several forms of PVDF crystal with the β form (form I) is the most reported [58,59,[66], [67], [68], [69]]. Other known crystal forms include α (form II) and γ (form III) [67,68,70] as well as δ (form IV) [66,71,72] and ε [71,73]. The structure of the α, β, and γ forms are shown in Fig. 2a. The structural details of the various crystal forms have been discussed in several publications [67,68,74,75]. Infrared spectra suggests that δ and α phase have similar chain conformation [76]. The α phase is thermodynamically the most stable under ambient conditions [72]. However, due to the anti-parallel packing of dipoles makes the α and ε phases nonpolar [73]. Only the α, β, and γ form display ferroelectric and piezoelectric properties. During melting of PVDF, several different peaks are observed [77] indicating the presence of several forms of crystals. The forms of crystal depend on the method of processing. For example, in the melt phase the α form is obtained by crystallizing below a temperature of 160^0^C, while at melt temperatures above this produces a mixture of α and γ phases. The γ phase content increases with increasing temperature [77]. The β phase is the most relevant in piezoelectric problems since it provides the best piezoelectric characteristics due zigzag conformation that can induce dipole moment. Several processing techniques can be used to obtain the β phase (Fig. 2b) as summarized by Ruan et al. [71]. The most common method to obtain the β phase can be obtained by mechanical drawing of the α phase [78] which aligns the dipole in a parallel configuration and generates the maximum dipole moment. Annealing at high temperatures transforms into γ phase with α and β as the minor phase [79], while annealing at high pressures converts it into β phase [68]. Application of high electric field can transform α-PVDF to a polar analogue i.e, the δ phase [66,80]. The molecular mechanism for the transformation has been proposed [81]. The process of electrospinning which results in uniaxial extension also results in the formation of the β phase [82].Fig. 2a) Diagrams of the α, β and γ polymorphic phases of PVDF. Reproduced with permission from Ref. [69]. b) Representation of the four methods of phase transitions to obtain β phase of PVDF. Reproduced with permission from Ref. [71]. c) Overview of the processing, properties and overview of barium titanate nanoparticles. Reproduced with permission from Ref. [154]. d) Photograph of wearable pressure sensor for heart rate monitoring fabricated using zinc oxide nanoneedles and PVDF hybrid film. Reproduced with permission from Ref. [71].Fig. 2
The crystal structure of PVDF films is significantly affected by the type of solvent used during casting [67]. For instance, casting PVDF sheets using acetone, hexamethylphosphoramide (HMPA), dimethylacetamide (DMAc) typically yields the α, β, and γ crystalline respectively. Interestingly, a single solvent (HMPA) can also be used to produce different crystalline form by using different solvent evaporation rate [83]. Several parameters such as solute concentration, solvent evaporation rate, temperature and pressure during evaporation have been known to affect the crystal structures of PVDF [[83], [84], [85], [86]]. The formation of α phase is generally favored at high solvent evaporation rate, while the slower evaporation rates favor the formation of β phase while intermediate rates produces a mixture of both phases [84]. In general, polar solvents induce polar β and γ phase upon solvent casting particularly when film thickness is in micrometer range. The extent of polar phase depends on the solvent evaporation rate [85]. Improved control of crystal structure are obtained by immersing the casting solvent solution in an antisolvent coagulation bath [87]. Additionally, incorporating small amounts of ionic liquid in solvents promote the formation of β phase in PVDF films [88] by promoting Coulombic interaction between the hydrogen atoms in PVDF and the anions in the ionic liquid. Short electric pulses can transform films into δ phase [72] which holds particular promise for usage in electronic industry (Fig. 2d). In addition to the factors discussed above, the crystal composition is also affected by environmental factors such as humidity and the nature of the substrate material on which the PVDF film is deposited [89]. These authors observed that higher humidity produces a higher fraction of γ crystalline phase in the cast film. While solvent cast films had a variable composition, solution cast films were more uniform.
Various processing techniques can be used to manipulate the crystalline phases of PVDF. During melt processing at high temperatures, two spherulite types were observed [[90], [91], [92]]. The most common is the α form characterized by their large sizes, while the smaller size spherulites have been identified as the γ form [93,94]. Subjecting this polymer to different mechanical treatment can produce different form of crystals [95]. The β phase can be obtained by drawing a polymer initially containing the α phase [[96], [97], [98]]. However, the processing history affects the formation of the β phase. It has been reported [99] that stretched injection molded samples resulted in the appearance of a new melting peak attributed to the β conformation, while in stretched compression molded samples this peak was absent. Additionally, cooling rate plays a significant role – with high or moderate cooling rate favoring the α form [77]. At low draw temperatures (<100^0^C), the α phase is transformed to the β phase occur at draw ratios of approximately 1. At higher temperatures the β phase is obtained only at draw ratios approaching approximately 6.7 [100]. For compression molded samples, the maximum content of the β form was obtained at 87^0^C [98]. At molding temperature around 160^0^C, discoloration of the polymer occurs [94] indicating the onset of thermal degradation. There is a critical stress which when exceeded uncoils the α phase to the planar zigzag β phase [101]. It should be noted that a residual α phase is present under all conditions. The γ phase can be obtained by melt crystallization at high temperature [90,91] and high pressure [74,102].
Similarly, poling also induces phase transitions in PVDF polymers. Poling aligns dipoles in the material upon the application of electric field. PVDF is one of the polymers that also display ferroelectricity [22]. By carefully controlling the magnitude of applied electric field one can control the phase formation of PVDF [64,66,80,[103], [104], [105], [106], [107]]. Under similar poling conditions, uniaxially drawn films had higher piezoelectric outputs [105]. Newman [103] reported that the phase formation depends on the strength of the poling conditions. At low fields (<1 MV/cm), the films were predominantly in the α phase and piezoelectric activity was small. In the range between (1 MV/cm < E < 2 MV/cm) there was a significant increase in piezoelectric activity with α phase crystallites becoming more polar. This is similar to what was reported by Davis et al. [66]. Further increase in the electric field to 3.2 MV/cm there was little change in piezoelectric values and no change in the structure and molecular orientation. When the electric field reached 4 MV/cm, a significant increase in the piezoelectric values were observed with the crystals transforming to the β form. Films containing different volume fractions of β and α phases, displayed a linear relationship between piezoelectric values and β fractions [107]. At higher electric fields the α phase can be converted to the polar δ phase [80].
Polyesters represent another class of material that display piezoelectric properties. Aliphatic polyesters are biodegradable and can be synthesized from natural resources (poly (lactic-acid), poly-β-hydroxybutyrate and copolymers and others). These are sometimes referred to as biological polymers. The presence of an asymmetric carbon atom on the polar atomic group of these polymers contributes towards their optical activity and their piezoelectric properties.
Among the polymers synthesized from natural sources poly(L lactic acid) (PLLA) is of interest. Depending on processing conditions two crystal forms (α and β) are observed [108,109]. The α-structure is observed at lower drawing temperature or lower draw ratio, while the β structure is observed at higher draw temperature or higher draw ratio. When crystallizing from dilute solution, PLLA results in the α form [110,111]. A third polymorph, known as the γ form is observed during epitaxial crystallization in hexamethylbenezene [112]. For interested readers, several studies on crystal structure of poly lactides are available [110,[112], [113], [114], [115], [116]].
Piezoelectric properties are comparable to that of PVDF [[117], [118], [119]] though they are much smaller than PZT ceramics. This has been attributed to the type of bond in the two materials – PLLA has covalent bond which requires higher force to displace than the ionic bond in ceramics that responds more readily to mechanical stress. Piezoelectric properties are imparted to the polyester by uniaxial stretching [37,120,121] without any required poling which is often challenging to apply in practice [122]. As a chiral polymer, PLA have carbon atoms with helical orientation. The two optical isomers of PLA are poly(L-lactic acid) and poly(D-lactic acid). When subjected to uniaxial extension display shear piezoelectricity [37], which arises from the rotation of polar atomic groups associated with asymmetric carbon atoms [118]. The magnitude of the piezoelectric constant is influenced both by the degree of crystallinity and molecular orientation. PLLA crystal has three non-zero piezoelectric constants (d14, d25, and d36) [120]. In a drawn PLLA film, the d14 and d25 are non-zero as the symmetry of PLLA films differs from that of the PLLA crystal [123]. Additionally, the chirality of the asymmetric carbon atoms determines the sign of shear piezoelectric constants.
The magnitude of shear piezoelectric constants correlate linearly with the product of polymer crystallinity and molecular orientation rather than with either of these parameters alone [36]. From a processing point of view, this is advantageous as achieving higher degree of orientation will inherently lead to higher crystallinity. Another technique used to improve the piezoelectric performance of PLLA film is to create a laminate alternating between PLLA and PDLA [117]. This lamination is possible because the piezoelectric constants of the two polymers have opposite signs [124]. Notable, a 120 multilayer film of PDLA/PLLA has a piezoelectric constant 20 times that of a single layer PLLA film. Further enhancement has been achieved by annealing PLLA films by treating with supercritical carbon-dioxide [125] which resulted in a significantly higher shear piezoelectric (d14) values than untreated films. For example, the application of this treatment to PPLA film with a MW of 200,000, induced a more condensed and uniform higher ordered structure, leading to higher piezoelectric coefficient. Another approach to increase the piezoelectric properties involves the addition of triblock copolymers to PLLA [126]. The addition of Poly(methyl methacrylate)-b-poly(butyl acrylate)-b- poly(methyl methacrylate) [PMMA-b-PBA-b-PMMA when added to PPLA doubled the piezoelectric coefficient by modifying the structure of PLLA/PMMA-b-PBA-PMMA blends. The PMMA end blocks at either end of the triblock copolymer is miscible with PLLA and acts as an anchor enhancing compatibility.
The piezoelectric response of PLLA can also be enhanced by changing the effective modulus and how stress is transferred to the piezo-active chains. Introducing rigid, well-bonded fillers (ceramics or mineral nucleating agents) raises the Young's modulus and crystallinity of PLLA and so improves force transmission to oriented crystallites. For example, diatomite added to electrospun PLLA fibers increased crystallinity, stiffness and device output (≈2.5 × vs neat PLLA) [127]. Embedding piezoelectric ceramic nanoparticles such as BaTiO3 into PLLA fibers and optimizing post-draw conditions also increases chain orientation, crystal phase fraction and electromechanical output [128]. Importantly, interface quality controls whether the added stiffness helps as surface functionalization, such as polydopamine coatings on BaTiO3, improves matrix–filler adhesion and force transfer, producing larger voltage outputs than uncoated particles at comparable loadings [129]. Another strategy is the engineering of morphology and thermal-strain, where careful combinations of drawing, thermal annealing and controlled crystallization increase both molecular orientation and crystalline modulus [127,130]. Architectural strategies, graded or multilayer stacks and engineered micro-geometries, such as aligned microfibers, perforations or porosity patterns, can concentrate shear/strain in piezo-active zones, effectively increasing the local mechanical impedance seen by polar groups without globally immobilizing chains [127,129].
PLLA also has an advantage over PVDF in the application of sensors. Notably, PPLA is not pyroelectric and hence, it does not generate unwanted voltage in response to temperature change. This makes it particularly well-suited for wearable sensors in contact with the human body, where heat generation is unavoidable. Another advantage is that in PLLA based sensor, the polarity of charge upon the application of tension is determined on the chirality and helical winding direction of PLLA. For PVDF, under tension the polarity is determined by the thickness direction of polarization [131].
Another optically active polyester synthesized from natural polymer is poly(β-hydroxybutyrate) (PHB) and its copolymers. PHB adopts a lefthanded helical structure with a twofold axis along the chain corresponding to two residues per turn in the crystalline state [39,132]. Piezoelectric relaxations corresponding to the three mechanical relaxations during temperature sweep was observed. The three temperatures at which relaxation processes occur are −120^0^C (γ-relaxation), 20^0^C (β-relaxation), and 130^0^C (α-relaxation) respectively. The room temperature piezoelectric relaxation (β-relaxation) is attributed to the increased piezoelectricity in the oriented non-crystalline phase. In copolymers of β-hydroxybutyrate and β-hydroxyvalerate (HV) the piezoelectric coefficients remain constant [38] suggesting that in the non-crystalline phase piezoelectric relaxations occurs due to relaxations in the elastic and dielectric mode. However, these copolymers exhibit reduced molecular stiffness and hence, reduced oriented non-crystalline phase.
Several other polymers are known to exhibit piezoelectric characteristics. These include polypropylene oxide [133], polypropylene [134,135], polyurethane [25,136], polyacrylonitrile (PANI) [137], polyethylene oxide [138], polyurea [[25], [26], [27],139], polyvinylchloride [140,141], and various elastomers [35]. Despite being largely amorphous, PVC exhibit piezoelectric behavior when subjected to rolling induced stress. The rolling process aligns dipoles along a preferred direction where the dipoles are distorted leading to polarization. In the case of PANI, polarization arises from two sources – a nonpolymeric species that carry charges (either transiently or permanently), and the second source arises from orientation of the nitrile side groups at temperatures above 90^0^C become thermally activated, contributing to both dielectric and mechanical relaxation [142,143]. Polypropylene, produces a large piezoelectric constant when processed by pressure expansion [144]. PP also displayed a decay in piezoelectric (e33 d33) values for a day before it stabilized [134]. Polyurethane displays transverse piezoelectricity which was caused by electrostriction [136]. Aromatic polyurea, on the other hand display a piezoelectric stress constant that is temperature invariant (−500C≤T≤2000C) while the strain constant increases owing to the decrease in modulus with increasing temperature [26]. For poled aromatic polyurea Piezoelectric constants is influenced by the ratio of the monomers used; the value being highest when the monomers are stoichiometrically equal [145,146].
The piezoelectricity of polymers is significantly lower, by nearly an order of magnitude-than that of ceramic materials (Table 2). However, ceramics are rigid and brittle, while polymers are easily processable and deformable. Combining the higher piezoelectricity of ceramics with superior processing capability of polymers is suitable for creating ‘hybrids’ with balanced properties and functionalities. As a result, composites of polymer and ceramic materials are an attractive candidate for fabricating piezoelectric composites [[147], [148], [149]]. Blending ceramics with polymers has the advantage of improving some of the properties such as piezoelectricity, brittleness of the inorganic ceramics, biocompatibility and ease of processability [150]. However, blending with polymers will affect the electrical properties of the composites by affecting such factors as the crystallinity and Coulombic potential [149]. Several review articles provide comprehensive discussion on the properties of piezoelectric composites [151] and the various methods of processing [71].Table 2Piezoelectric properties of selected synthetic polymers.Table 2MaterialsPoling Field (kV/cm)Piezoelectric constant (pC/N)CommentsReferencePANI500.65d31@25^0^C(11 Hz)[137]PVDF500Up to 23d31[537]500Up to 33d33[537]Nylon 1110003–14d31[537]10002d33[537]9004.2d31[32]Nylon 91.1d31[277]Nylon 5/70.4d31[33]Polyurea10d31[25]BaTiO34.578d31[537]Up to 100d33[537]PVF2001.32–1.00d31[21]PVC1800.5–0.33d31[21]PC2000.1d31[21]
Quartz possesses a crystal structure that exhibit natural piezoelectricity [152]. Other well-known ceramics with piezoelectric properties include lead-zirconate-titanate {(Pb[ZrxTi1-x]O3) 0≤x≤1}, Barium Titanate (BaTiO3), Sodium potassium niobate [(Na, K)NbO3], and Zinc Oxide (ZnO). These ceramics have been applied as fillers to enhance piezoelectricity of polymers. Many of these exhibit a perovskite crystal structure, with a general structure of ABO3 where A and B are cations. In this structure, A represents a lanthanide or an alkali earth metal while B represents a transition metal [153]. Because of its excellent ferroelectric properties, dielectric constant and biocompatibility, BaTiO3 has been widely studied as filler in bio-piezoelectric composites [154]. Its non-centrosymmetric tetragonal crystal structure enables the material to display piezoelectric behavior. The size of the particle also plays an important role in the properties of the composites [155]. As the grain size decreases, the dielectric properties improve [154,156] and resulted in enhanced stability of the crystal phase at room temperature. Piezoelectric coefficients of fibers of BaTiO3 dispersed in poly(vinylpyrrolidone) can be improved by decreasing the fiber diameter [157]. Various morphology and size of nanoscale nonconducting structures have been produced and include nanowires [158,159], nanoparticles [160,161], nanorods [162], nanocubes [163] among others [154]. Shape and size of fillers also the affect the dielectric properties of composites [164,165]. Notably, the dielectric constant tends to increase as the nanowire diameter is reduced below a critical threshold [166]. Beyond this critical diameter the ferroelectric behavior approaches that of the bulk materials.
Nanoparticles of BaTiO3 can be synthesized using various methods which can be categorized as catecholates [167], hydrothermal [159,[168], [169], [170]], solvothermal [171], and co-precipitation [172]. These particles have been blended with polymers such as PVDF [148,[173], [174], [175], [176]], PLA [177,178], cellulose [179], chitosan [180] and others [148,154].
However, to achieve uniform dispersion for BaTiO3 fillers within the polymer matrix, its surface needs to be modified. Due to its large surface area per unit mass and high surface energy, BaTiO3 nanoparticles tend to form aggregates in aqueous media, leading to a decline in properties as BaTiO3 loading increased [174]. Various techniques are available for functionalizing the surfaces of BaTiO3 nanoparticles [154,181,182] to mitigate aggregation and enhance compatibility with polymer matrices. These including hydroxylation of surface by thermal decomposition of hydrogen peroxide [[183], [184], [185], [186]], bifunctional silane coupling [[187], [188], [189], [190]], amine [191,192] and carboxyl functionalization surfaces [193], dopamine functionalization [174,[194], [195], [196]] and grafting with polyethylene glycol [197,198]. BaTiO3 NPs can be functionalized with polystyrene to improve its compatibility with PS-b-PMMA block copolymer [199]. The negative charge on the surface of BaTiO3 NPs can be utilized to attach to the positively charged surfaces [179]. In each instance the piezoelectric properties increased significantly over mixing with unmodified BaTiO3. For example, the polarization coefficient (d33) of pure paper increased 12-fold upon the addition of 48 % BaTiO3 [179]. A small fraction (1 % by volume) loading of dopamine modified BaTiO3 NP increased the dielectric breakdown strength of PVDF nanocomposites by 150 % [174]. It was speculated that the dopamine acts as a molecular bridge between the BaTiO3 NP and PVDF. Calcination of BaTiO3 at high temperatures (>850^0^C) improved their properties over uncalcinated ceramics in BaTiO3/PVDF mixtures [175] (Fig. 2c). Similarly, the introduction of gadolinium doped BaTiO3 into poly(lactic-co-glycolic acid) caused a significant improvement in the composites dielectric and piezoelectric properties [200]. Plasma enhanced chemical vapor deposition (PECVD) introduced amine functionalities on BaTiO3 surface [201], resulting in a more uniform distribution of the nanoparticle within an epoxide matrix over unmodified particles resulting in better properties at each level of loading.
Conductive fillers like graphene [[202], [203], [204], [205], [206], [207], [208], [209], [210], [211]], nanotubes [[212], [213], [214], [215], [216]], nanowires and carbon black [217] have been utilized to achieve higher dielectric constants at a relatively low filler volume compared to nonconducting fillers. There is, however, a critical concentration known as percolation threshold beyond which there is a sharp rise in the electrical conductivity causing the composite to become conducting [[218], [219], [220], [221]]. Ram et al. [221] reported that the percolation limit is controlled by the aspect ratio of the filler. The incorporation of a small amount (∼0.1 %) reduced graphene oxide (rGO) led to a significant piezoelectric and ferroelectric properties of PVDF/rGO composites [204,211]. Alamusi et al. [202] reported that piezoelectric PVDF/rGO composites was approximately three times higher than that of pure PVDF. Similarly, the addition of 0.1 % graphene oxide (GO) increased the β-polymorph in PVDF leading to increased piezoelectric properties [207]. Incorporation of polyaniline functionalized GO to PVDF-TrFE bilayer films also led to a notable enhancement in properties [210]. 3D printed PVDF containing either graphene or multiwalled CNT demonstrated increased β-phase and crystallinity and thus, resulting in enhancement of their piezoelectric properties [222]. In these composites, the graphene and MWCNT acted as a nucleating agent, with the graphene being more efficient because of its lamellar structure. Additionally, 3D printed MWCT/BaTiO3/PVDF composite produced homogeneous films with enhanced dielectric constants [223].
Microwave processed multiple walled carbon nanotubes (MWCNT) embedded in epoxy samples had a dielectric constant that 2.5 times larger than the traditional thermally processed composites [214]. Sanati et al. [215] developed a CNT/PVDF piezoelectric composite sensor whose sensitivity exceeded that of commercially available sensors. Furthermore, the addition of 0.1 % by weight of CNT enhanced the performance of PVDF/potassium sodium niobate (KNN) electrospun composites by increasing the amount of β-polymorph in the hybrid composite material [216].
The properties of composites depend significantly on the state and connectivity of the constituent phases contained. In diphasic system, an initial classification identified 10 possible connectivity types [224,225], which was later expanded to include six additional phases [226]. Each phase can be continuous in 0, 1, 2, 3 dimensions. Hence, there are 4^2^ possible connectivity combinations. These configurations are denoted as (m-n), where m refers to the dimensional connectivity of the piezoelectric (active) phase and n represents the dimensional connectivity of the polymeric (passive) phase. These combinations are (0-0), (0–1), (0–2), (0–3), (1-0), (1-1), (1–2), (1–3), (2-0), (2-1), (2-2), (2–3), (3-0), (3-1), (3-2), (3-3). The first digit represents the number of dimensions of the piezoelectric active phase and the second digit is used for the polymeric phase. For example, (0–3) combination represents a particle dispersed in a continuous polymeric phase. Comprehensive reviews of various techniques used in developing piezoelectric composites and exploring these connectivity types are available in the literature [151,[227], [228], [229], [230], [231]].
One effective strategy to increase the piezoelectric properties of polymers is through controlling the crystallinity. This can be achieved by processing a polymer above its melting point, extruded it, and collecting it using a rotating drum which promotes alignment and crystallization [232]. In one example, a P(VDF-TrFE) piezoelectric polymer layer sandwiched between conductive polycarbonate electrode layers produced a fiber like structure. Coaxial melt spinning is another processing technique, in which a piezoelectric sheath surrounds a core material. The core can be a conductive metal [131], a polymer [233], or a polymer containing conducting filler [234,235]. The incorporation of fillers such as oxides and CNT's can increase the piezoelectricity by acting as nucleating agents that promote crystallization due to the formation of polar β-phase. For instance, the addition of 0.01 % by weight of amine functionalized CNT's increased the β-phase of cold drawn PVDF fiber [236]. Similarly, the addition of 5 % by wt clay and increasing the draw ratio also increased the β-phase of PVDF polymer [237].
Most of the composites that are useful consists of ceramics dispersed in a polymeric phase, with the primary goal of increasing the piezoelectric response of the polymer. These fillers take the form of nanoparticles, nanorods, or nano-platelets dispersed in a continuous polymeric phase. Hence, one could envision a (0–3), (1–3) or a (2–3) connectivity composite being fabricated.
Several fabrication techniques are available to produce such composites including melt processing (extrusion, injection molding), solvent casting, 3D printing and electrospinning. The (0–3) pattern is easy to fabricate and scale up [238]. For example, Banno and Saito [237,[239], [240], [241]] produced blends of piezoceramics and synthetic rubber using hot roller to produce 0.5 mm thick sheets. Similarly, lead zirconate titanate (PZT) dispersed in polydimethylsiloxane (PDMS) were used to fabricate thin sheets (130–160 μm thick) via spin coating [242]. A (0–3) composite consisting of a piezeoceramic and polyurethane was obtained using solvent casting [243]. Nanorods of gallium wrapped PZT was dispersed in P(VDF-TrFE) were also developed using solvent casting method [244]. Other fabrication techniques include dispersing CNT in silk fiber suspension and sonicated to achieve uniform dispersion and subjecting the suspension to vacuum filtration until completely dry solid composite [245]. PVDF/BaTiO3 in the required ratio were melt compounded using an extruder and the extrudate was cryogenically ground, mixed with fumed silica to improve flowability before being shaped using selective laser sintering process [246]. Fiber mats consisting of BaTiO3/PVDF were prepared using electrospinning [247]. Lu et al. [248] developed composites of BaTiO3/PVDF, PZT/PVDF, and CNT/PVDF composite by fiber drawing. In addition to melt spinning, solution spinning has also been attempted [249,250]. In this method, the polymer is dissolved in a solvent and extruded through a spinneret. The extruded solution is collected in a coagulation bath forming the fibers. The residual solvent is subsequently removed through additional washing steps.
Biological macromolecules such as cellulose, collagen, DNA and proteins display piezoelectric behavior. Piezoelectricity has been detected in bone [251], tendons [252,253] (Table 3), silk fibers [50] fibrous proteins such as keratin [254], myosin and actin [255], elastin [256] and wood [257,258] The piezoelectric coefficient in these materials is dominated by the shear components of piezoelectricity – where the electric polarization is produced in the direction that is perpendicular to the plane of shearing stress. This is characteristic of optically active polymers where the contribution of tensile or compressive stresses is a minimum. Piezoelectricity in these biopolymers are derived from the internal rotation of polar atomic groups attached to the asymmetric carbon atoms [43] and can be generally attributed to the organic component in the material – collagen for biological tissues and cellulose for wood [16,259].Table 3Piezoelectric constants of systems of structural proteins (pC/N). Unit: 10^−8^ c.g.s.e.s.u. Adapted with permission from Fukada E. and Yasuda I. 1964, Jpn. J. Appl. Phys. 3 (1964) 502B.Table 3Materiald14d15d33d31Achilles Tendon (Bovine)−2.61.40.070.09Achilles Tendon (Horse)−1.90.520.070.01Femur (Horse)−0.220.040.0030.003Silk−1.10.230.020.02Wool−0.070.070.0030.01
Piezoelectric properties have been investigated for DNA [260], polypeptides [52,53,[261], [262], [263], [264]] and even sea shells [53,265,266]. DNA is a molecule that is present in most human cells and nearly all other organisms, specifically located in the cell nucleus. Ando and Fukada [260] proposed that the origin of the polarization in DNA is induced by the reorientation of dipoles in sugar-phosphate backbone. Additionally, the polarization resulting from the stress-induced orientation of the bases is believed to contribute an effect of opposite sign to that of the backbone chain.
DNA, the main building blocks of life, encodes the instruction that determines the sequence of amino acids in proteins. The majority of the 20 known amino acids that make up protein structure do not form centrosymmetric crystal structures [267,268]. As a result the majority of protein structures are non-centrosymmetric which makes them piezoelectric [269]. Vasilescu et al. [270] reported that 19 out of the 20 amino acids exhibited either theoretical and/or experimental piezoelectricity. The physical origin of piezoelectricity in amino acids lies in their molecular asymmetry and the symmetry properties of their crystalline arrangements. Amino acids are chiral molecules (except glycine), meaning they lack inversion symmetry at the molecular level. When these molecules assemble into crystals, their packing often reinforces this lack of centrosymmetry. Since piezoelectricity requires a crystal structure without a center of symmetry, amino acids are natural candidates for exhibiting this property. In fact, the piezoelectric response emerges because mechanical stress perturbs the asymmetric charge distribution of the molecular dipoles, generating a net polarization. Furthermore, in proteins, the hierarchical organization of amino acids into secondary (α-helices, β-sheets) and tertiary structures amplifies these effects. For example, α-helical domains, rich in aligned peptide dipoles, create a cumulative polarization that can give rise to pronounced piezoelectric responses at the macromolecular scale. Thus, the piezoelectric behavior of proteins is rooted both in the intrinsic chirality and polarity of their amino acid building blocks and in the larger non-centrosymmetric arrangements that arise from protein folding and supramolecular organization [267]. Under electrical stimulation, γ-glycine, L-arginine and DL-alanine have demonstrated the highest piezoelectric activity surpassing even that of quartz crystals [271,272]. In contrast, L-alanine, L-valine, L-glutamic acid and DL-tyrosine exhibited much smaller piezoelectric effect (Table 4). A possible explanation of this variation could be attributed to the crystallinity or the inherent differences between the optical activity that proteins possess. Additionally, a subsequent study found that the piezoelectric activity of amino acid crystals is temperature-dependent [272]. This temperature effect was not observed in all samples, which was attributed to enhanced damping of elastic vibrations in the crystals due to rotation of the CH3 and NH3 groups.Table 4Piezoelectric activity of different amino acids based in theoretical and experimental observations.Table 4Amino acidTheoretical PiezoelectricityExperimental Piezoelectricity by Vasilescu (Vasilescu, 1970)Experimental Piezoelectricity by Lemanov (Lemanov, 2000)Symmetryα-glycine-^a^–+C2hγ-glycine+^a^++C3L-alanine+^a^++D2L-valine+^a^–+C2L-isoleucine+^b^+–C2, D2L-glutamic acid+^c^–+C2, D2L-cysteine+^d^+–C4, D6DL-alanine+^c^++C2vDL-valine+/−^c^––C1DL-serine+^e^+–C2hDL-aspartic acid-^f^––C2hDL-lysine+^g^–D2DL-tyrosine+^h^++C2vDL- tryptophan+^i^––C1References of the theoretical piezoelectricity [[538], [539], [540], [541], [542], [543], [544], [545], [546]].
In living systems, proteins perform multifaceted functions including a range of structural and catalytic properties due to its varying amino acid composition. It has been proposed that piezoelectricity could also be involved in enzyme activity [273] given the inherent piezoelectric properties of amino acids. Electrical polarization of the surrounding medium may arise from ions or charged species – such as those produced as a result, of chemical reaction - which become trapped at some localized enzyme sites. A theoretical model has been proposed in which enzymes are treated as piezoelectric semiconductors and substrates are considered as capable of electrically activating the corresponding enzymes [274].
For polypeptides such as poly-γ-methyl-L-glutamate and poly-γ-benzyl-L-glutamate, exhibit piezoelectric properties both α-helix and the β-form conformation [52,261,263]. Similarly, poly-β-benzyl-L-aspartate which also display an ω-helix molecular conformation also displays piezoelectricity [18]. In addition, an additional random coil conformation is not piezoelectric because of the presence of center of symmetry. The electrical polarization observed under shear stress is attributed to the internal rotation of the dipolar groups around the asymmetric carbon atoms in the polypeptide chain [275]. Poly-γ-benzyl-L-glutamate can also be induced to exhibit piezoelectric properties through solvent casting under strong magnetic field which orients the sheet [276]. It has been reported that poly-γ-benzyl-L-glutamate forms liquid crystal phase in solvent [118]. In a related, study, poly-α-isobutyl-L–aspartate in chloroform solution evaporated under strong electric field produced highly piezoelectric films [277]. This polypeptide has a α-helix like rigid helical formation found in polypeptides. The oriented film has a structure like the liquid crystalline phase.
Oriented DNA samples cast from ethanol solution and stretched to 1.6 times its original length also displayed piezoelectric [260]. This was attributed to the reorientation of the polar groups within the DNA molecules under applied stress. The effect of moisture content on the piezoelectric properties of DNA and cross-linked poly(hydroxyethyl glutamine)(PHEG) has also been investigated. The piezoelectric constants were found to decrease with increased adsorbed water while complete dehydration reverses the sign of the piezoelectric constant [278]. The phenomena were attributed to the disordered state of DNA upon dehydration. For the polypeptide PHEG, the increase in water content was found to increase the dielectric constant [275].
High piezoelectric properties have been observed in films and bulk crystals of orthocarboranyl derivatives of (S)-asparagine and (S) glutamine [279]. Piezoelectric behavior has also been observed in the globular protein lysosome [280], bioinspired protein nanotubes [264], and diphenylalanine peptide [281]. Seashells - composed of calcium carbonate (95 %), biopolymers and a small amount of water have been found to exhibit both piezoelectric and ferroelectric properties [54]. The biopolymers in nacre consists of chitin and various proteins such as perlucin, perlwapin, AP8, and perlinhibin [282,283]. The most prominent amino acids are aspartic acid and asparagine, glutamic acid and glutamine, proline, glycine and leucine [284]. With the exception of L-glutamine, all chiral protein amino acid crystal exhibit piezoelectricity [272] with c-glycine having the greatest piezoelectric response. In nacre, the piezoresponse is directional with the intercrystalline biopolymers displaying strong response along the direction perpendicular to the mineral platelets [54]. The interlamellar biopolymer provides a strong response in a direction parallel to the platelets.
The role that electrical stimulus has in facilitating osteoinduction has been long recognized [49]. The development and remodeling of bone are influenced by its piezoelectric properties [16,251,[285], [286], [287]], suggesting the existence of a relationship between the electrical polarization and the resulting stresses encountered in the bone and its internal structure. Bone has two principal crystalline components – hydroxyapatite (∼65 % by mass) which is centrosymmetric and is neither piezoelectric or pyroelectric, and collagen molecule (∼35 % by mass) that can exhibit both. For example, tooth enamel which is composed of hydroxyapatite is characterized by the absence of piezoelectricity, but present in dentine mainly composed of mineralized connective tissues [288]. In collagen, the polypeptide chains are arranged in a coiled-coil triple helix forming a long rod-like molecule [289]. The polar crystal structure of collagen in the direction of bone axis results in bone being piezoelectric [252,253,290]. Indeed, as early as 1957 piezoelectric effects were observed in the Achilles tendon of horse and cow [251] (Table 3). The high degree of orientation in collagen makes it conducive to the piezoelectricity. The high elastic moduli of hydroxyapatite crystals enables load placed on the bone to be transmitted to the more compliant collagen (which has a lower moduli), inducing the deformation needed for generating piezoelectricity effects [40,252]. Experimental mechanism supports this mechanism that decollagenation of bone eliminates the piezoelectric effect, while demineralization does not [290]. This evidence was originally verified by Fukada and Yasuda [251,253]. The authors observed a direct relationship between polarization and stress, assuming that the effect was truly piezoelectric, which results from the slipping of collagen fibers past one to another. It has also been proposed, that deformation of hydrogen bonds or its cross-linkages within collagen under stress induce change in the spontaneous polarization of collagen [252]. Furthermore, models of bone remodeling has been developed and compared with experimental and clinical data [46,47]. The results indicate that these piezoelectric models can successfully predict various aspects of bone remodeling.
In the 1960's collagen piezoelectricity was the primary factor responsible for providing stimulus to bone cells. However, bone's piezoelectricity is not just confined to the collagen alone – not even to the stoichiometric HAp. Piezoelectric activity was detected on HAp ceramics at a macroscopic scale [291]. While the apatite crystals in bone closely resembles HAp, they contain impurities, which give rise to a disordered arrangement of the hydroxyl (OH^−^) ions [292]. In this sense, the structural protein, collagen, as well as the inorganic crystals of HAp, when considered independently, exhibit electrical phenomena - albeit in different piezoelectric forms. Marino and Becker [293] have proposed the junction of these molecules as the relevant area of the piezoelectric effect production. They further suggested that cross-linkages between molecules may serve as important sites for charge production.
The moisture content of bone significantly influences its piezoelectric properties [42,44,251,294]. While dry bone [16,[251], [252], [253],287,295,296] has been documented to exhibit piezoelectricity the same cannot be conclusively stated for wet bone. There is also a notable difference in the properties of dry and wet bone, especially in their dielectric characteristics [297]. Dried bone display higher piezoelectric than wet collagen [251,298] which contradicts the results from Anderson and Eriksson [44] who observed increasing values of shear piezoelectric coefficient for wet bone. Gaining a thorough understanding of piezoelectric effects in bone under physiological conditions is important if their mechanism and function is to be elucidated. While mechanically stressed bone generates electrical potential, the origin of the stress-generated potential (SGP) (piezoelectricity and streaming potential) under biological conditions remains a subject of debate. Hence, the findings of Anderson and Erickson [42,44], which proposes both piezoelectricity and streaming potential contribute to SGP has some validity. Subsequent studies [[299], [300], [301], [302]] have further highlighted the importance of SGP in bone remodeling process.
The negatively charged surface of bone matrix interacts with bone extracellular fluid. At the slip plane, which is the interface that separates the fluid attached to the surface and the mobile surrounding fluid, an electric (zeta) potential is generated. The SGP's are directly proportional to the zeta potential [302]. Experimental evidence [301] have demonstrated that the electrical potential in bone and tendon is produced by electrolyte flow. The presence of an electric potential without mechanical deformation of the tissue would cast doubt on the phenomena of piezoelectricity as the sole reason for this. The zeta potential changes sign at pH ∼4.3 for tendon [42]. Streaming potential displayed an inverse dependence on NaCl concentration, the solution viscosity, and calcium ion concentration in the solution [301]. These observations imply that streaming potential is the dominant effect in wet bone. Furthermore, studies have confirmed [303] when collagen in the bone is removed, the zeta potential decreases significantly. This is not to suggest that piezoelectricity is absent from wet bone. Using Piezoresponse Force Microscopy, Halperin [48] observed similar piezoelectric response between the dry and wet bone. Interestingly, when hydrated collagen sheets are frozen, thereby reducing the conductivity of ice compared to that of the water, they still exhibit piezoelectricity at −25^0^C [304]. However, the piezoelectric values of the frozen samples are significantly lower than their dry counterparts. In dental hard tissues, piezoelectric effect has been detected in dentine but not mineral apatite or enamel [305].
The primary amino acids that constitute collagen consist of proline, glycine, and hydroxyproline. At room temperature the simplest amino acid glycine can crystallize into the three polymorph α [306], β [307], and γ [308]. Since α form crystallizes in the centrosymmetric phase it is not piezoelectric. However, when doped with alanine, threonine, or serine reduces symmetry and induces polarization and piezoelectricity [309]. The β-polymorph of glycine is reported to be ferroelectric [268]. Using density functional theory (DFT) Guerin et al. [267] observed that the longitudinal piezoelectric coefficients for both γ and β polymorph were higher than the shear piezoelectric coefficient. Furthermore, these authors also reported that the magnitude of the longitudinal and shear components of piezoelectric coefficient for the β polymorph exceeds that of the γ-polymorph. Under ambient conditions the γ-form is the most stable, but transfers to the α-form at elevated temperature [308,310]. These authors were able to produce solution grown β-glycine which has several charged domain walls. An external bias can be used to control the charged domain walls which in turn affects the properties of β-glycine. Electrospinning technique has been employed to incorporate β-polymorph of glycine incorporated into poly (vinyl alcohol) [311], resulting in nanofibers where the crystal aligns along the axis and displayed enhanced piezoelectric properties. Under ambient conditions the β polymorph converts to the α form. The time taken for the transformation from the β to α form is retarded in nanoscale sized glycine crystal [312,313]. γ-glycine can also be synthesized from solution [314] and is both thermally stable and piezoelectric.
Single type-I collagen fibril predominantly display shear piezoelectric coefficients [253,315], a behavior which can be explained by its semi-hexagonal structure [253]. Similar observations were obtained for type II collagen fibril which also displayed shear piezoelectricity [316]. The values obtained were lower than the type I collagen, attributed to the variation in their structural organization. However, it was reported that rat tail tendons display both shear and longitudinal piezoelectric responses [317]. Similarly, cross-linking of collagen films induces vertical piezoelectricity [41]. Beyond collagen, elastin, a key protein component in the extra-cellular matrix (ECM), has also been reported to be ferroelectric behavior [318].
Dental hard tissues are chemically similar to bone and include three main hard enamel, dentin and cementum [319]. Branden and co-workers observed that compressive stimuli generated electrical potentials. However, using elephant dentin they demonstrated that enamel do not possess this property [305]. Later, Marino and Gross [288] studied the same phenomenon using whale teeth and human tibias. Their experimental data indicated that dentin and cementum have essentially nearly equal piezoelectric constants and they were able to produce approximately 12 % of the surface charge density produced by cortical bone under similar mechanical load conditions [288]. A different research group observed that the electromechanical constants obtained using fresh human molars were higher than those observed in whale teeth [320]. They also reported that piezoelectric properties of dentin were significantly affected by both humidity and orientation of the tubules. However, unlike bone, tooth itself does not exhibit a growth response to orthodontic force, because its biopiezoelectric effect is feeble.
Just as mechanical deformation induces electrical potential which has been linked with cellular signaling and bone remodeling, researchers have suggested that similar effects may occur in tendons [293,321]. Tendons are predominantly composed of cells, collagen fibers, proteoglycans and water. The collagen present in the connective tissues is mainly type I, although it contains small amounts of types III and V [322]. Fukada and Yasuda [253] supported this concept by detecting significant piezoelectric effect in dried Achilles tendon. Nevertheless, to accurately measure the mechanical behavior of tendons in vivo remains particularly challenging as the current methods have some limitations. Furthermore, a comparison of the results from Marino and Fukada's indicates that bone collagen exhibits one-sixth of the piezoelectric response observed in tendon collagen. This may be attributed to the highly orderly parallel organization of the protein fibrils in tendon or due to differences in the bonding of collagen molecules in each tissue type [251,253,290].
A tendon is a resilient band of fibrous connective tissue that connects muscle to bone - or muscle to muscle and is specifically designed to withstand tension. The piezoelectric phenomenon was observed in skeletal muscles, notably in rabbit psoas muscle [255,323]. Skeletal muscle is mainly composed of two different fibrous proteins, myosin and actin which exhibit shear piezoelectricity due the internal rotation of polar atomic groups associated with asymmetric carbon atoms [323]. Later studies have shown that films made up of myosin and actin, manifests piezoelectricity, though the magnitude of piezoelectric constant is significantly smaller than that observed in dried skeletal muscle. The researchers attributed the difference to a higher degree of molecular orientation and crystallization in intact muscles of the tissue [255].
Piezoelectric behavior has also been observed in smooth muscle, specifically in large blood vessel (aorta and vena) walls [324]. In aorta, these researchers verified that the physical property is similar to that found in bone, and this property depending on the direction of vessel elongation (the fibers in the longitudinal direction are responsible for this feature). Similar conclusions were observed for vena which depends on the degree of orientation and humidity content [324]. Furthermore, both tissues were shown to generate an electric current in the order of picoamperes due to the cyclic deformation of the blood vessel walls caused by blood vessel pulsation. Although such small current are unlikely to have influence on normal physiological functions, they may become significant in the event of tissue damage [324]. Under these conditions, the piezoelectrically induced current might play a key role in restoring normal physiological function. In fact, Shamos and Lavine [16] supposed that one aspect of this property in blood vessel would be the formation of thrombi following injury of blood vessels.
Fukada and Hara [324] also examined the piezoelectric effect in other types of tissues, namely trachea, intestines and ligament. The trachea and intestines used were taken from dog. The anisotropy observed in these tissues were similar to that in bone and tendon, and it was suggested that the piezoelectric fibers in intestine were aligned in a tubular direction, while those in trachea are arranged in a circumferential orientation [251]. These findings are consistent by their anatomical structure, since trachea is composed by circular cartilages which have collagen molecules oriented circumferentially [324]. To further explore the piezoelectric properties of elastin, the researchers extracted ligaments from bovine leg; the electric currents observed were comparable to those measured in bovine femur. They concluded that the orientation of elastin fibers lies along the fiber direction.
In the nervous system, a complementary body of experimental and theoretical work suggests that electrical action potentials are accompanied by measurable mechanical and volumetric changes, and that mechanical perturbations of membranes can produce electrical responses. Classic experimental studies using sensitive mechanical probes and optical/pressure measures demonstrated transient surface displacements and swelling coinciding with action potentials in excitable preparations. For example, Iwasa, Tasaki, and Gibbons reported small but reproducible outward surface displacements (tens of angstroms) and swelling pressure changes associated with propagated action potentials in crab and squid axons. The timing of these mechanical events matched electrophysiological signals, supporting a strong association between electrical activity and mechanical deformation in axons [325,326]. Building on this, nanometer-scale mechanical “spikes” have been documented in mammalian nerve terminals using high-bandwidth atomic-force microscopy. These mechanical events were temporally linked to the action potential and, based on concurrent optical and ionic measurements, were attributed to rapid, ion-flux–driven water movements that alter terminal volume. Direct AFM traces, together with complementary electrophysiological and optical data, confirmed that these mechanical changes are not artifacts but physiologically coupled to electrical impulses [327]. The experimental evidence has inspired theoretical treatments of electromechanical coupling in nerve membranes. Gross and colleagues formulated a model in which membrane surface charges, electrostriction/piezoelectric-like responses, and stress-dependent changes in membrane fields mediate mechanical-to-electrical and electrical-to-mechanical transduction. Their quantitative estimates predict small but physiologically plausible changes in axon geometry arising from membrane voltage changes. While such work does not establish that nerves exhibit classical crystal-type piezoelectricity (as in ferroelectrics), it formalizes how electrostatic, electrostrictive, and piezoelectric-like effects in a membrane–cytoskeleton system could couple mechanics to membrane potential, consistent with the measured amplitudes and timing of mechanical events [328].
At the molecular level, the piezoelectric phenomenon in biological systems arises from the non-centrosymmetric arrangement of proteins such as collagen, keratin, and chitin, which generate electric charges under mechanical stress. Within the nervous system, piezoelectricity is thought to emerge through mechanocoupling linkages involving voltage-gated ion channels and the oriented fibrous structures of the extracellular matrix surrounding nerve fibers and axons. This electromechanical coupling may contribute to sensory transduction and signal propagation in neural tissues [329]. Foundational experiments on biological proteins confirmed piezoelectric effects across both soft and hard tissues, including collagenous components resembling the perineurium and endoneurium in peripheral nerves. Stress-induced potentials were measured in these structures using electrometers on dried samples, supporting the hypothesis that piezoelectricity is a universal property of living matter [16]. Specific to nerve receptors, both piezoelectric and pyroelectric mechanisms have been proposed as the biophysical basis for detecting mechanical force and temperature changes. Experimental observations in stretch receptors, such as those of crayfish muscle spindles, revealed a logarithmic relationship between neural firing rates and applied force. This relationship aligns with piezoelectric theory, where charge generation is proportional to stress, and was quantified through electrophysiological recordings. These findings provide empirical support for piezoelectric transduction in neural mechanoreception without requiring additional energy sources [330].
The origin of the intrinsic mechanoelectrical conversion properties of skin tissues is thought to arise from electrokinetic phenomena in the dermis. The skin constitutes approximately 6 % of our total body weight. It incorporates all of the key support systems of our blood, muscle as well as innervations – and exhibits a complex biochemistry. Roughly 90 % of total dermal protein consists of collagen, (predominantly type I, III and V). The other major ECM fibrous protein is elastin, which is responsible for skin elasticity. Keratin is another protein that belongs to the skin's biochemical structure. This protein is produced by keratinocytes, which form the basis of epidermis [331]. Because these molecules possess a crystalline structure of an asymmetric nature, they exhibit electromechanical behavior [[9], [316], [317], [332]].
The first evidence of piezoelectricity in skin was reported by Shamos and Lavine [16], who perceived a direct piezoelectric effect in dry, whole-skinned samples from human forearm, as well as from cat, and pig specimens including callused specimen from the sole. They attributed this physical property with the oriented dermal collagen fibrils. These findings were further corroborated by examining human skin in vitro and in vivo which led to the conclusion that the electric signals obtained had piezoelectric nature, dependent of the tissues living state [334]. Moreover, these authors suggested that this piezoelectric property has its origin in uniaxial oriented epidermal keratin filaments in the basal cells layer. Origin of human skin piezoelectricity was examined from different skin sources such as breast, thigh, prepuce, and callus from sole and by separating the samples into their dermal, epidermal and horny layer components [335]. In agreement with the previous studies, it was reported that human dermis, true epidermis, and horny layer, all display biopiezoelectric activity. In accordance with prior observations [16] it was concluded that collagen network is responsible for this piezoelectric property in the upper layer [335]. Athenstaedt's further reinforced these findings, by showing that true epidermis exhibits the same piezoelectric property [334]. This effect was attributed to α-helical keratin-like fibrils, which accounts for the significant piezoelectric property observed in the horny layer [333,335].
Understanding the interplay between physiologically generated electric fields and mechanical properties at the molecular, cellular, and tissue levels has become the main motivation for exploiting piezoelectricity in biological systems through tissue engineering (TE) strategies. Many native tissues including bone, cartilage, skin, muscle, and nerves exhibit intrinsic piezoelectric properties that play essential roles in their development, remodeling, and repair. For example, collagen-rich tissues generate electric potentials when mechanically deformed, guiding cellular alignment and extracellular matrix organization, while excitable tissues such as nerves and muscles rely heavily on electrical cues for signal propagation and functional regeneration. These naturally occurring bioelectrical phenomena highlight a fundamental connection between mechanical forces, electrical activity, and biological outcomes.
However, when injury or disease disrupts the structure of a tissue, its intrinsic piezoelectric capacity is also compromised, leading to impaired self-repair. This limitation can be overcome by introducing piezoelectric materials into TE approaches. By embedding piezoelectric scaffolds into the defect site, it becomes possible to artificially restore or even enhance the natural electromechanical environment. Such scaffolds can convert physiological mechanical loads into localized electrical cues, thereby mimicking the native signals that orchestrate cell behavior, matrix deposition, and overall tissue regeneration. In this way, the piezoelectric effect acts not only as a biophysical stimulus but also as a bioinspired design principle for next-generation regenerative materials. Different tissues respond preferentially to distinct forms of electromechanical stimulation, mechanical loading and deformation are particularly critical for bone and cartilage, where piezoelectric scaffolds can amplify functional stress-induced signals, while controlled electric fields are indispensable for muscle and neural tissues, where they promote excitability, polarization, and synaptic integration.
Furthermore, bioelectrical signals are naturally generated by ion channels, transporters, and pumps within cells, and the spatial organization of these components, such as the apical-basal polarity of epithelial cells, creates a transepithelial potential that coordinates tissue-level responses [336]. This intrinsic electrogenic property means that virtually all cells, not only excitable ones, are sensitive to electrical cues and can use them to mount organized regenerative responses [337]. Building on this principle, piezoelectric scaffolds serve as a bridge between natural bioelectric signaling and engineered tissue constructs, offering a self-sustaining method to stimulate regeneration without the need for external electrodes or continuous electrical devices. Previous studies have demonstrated that a wide range of biological systems respond robustly to both endogenous and exogenous electric stimuli, reinforcing the concept that physiologically relevant electric fields can regulate cell proliferation, migration, differentiation, and tissue homeostasis [338]. Consequently, two-dimensional (2D) and three-dimensional (3D) scaffolds composed of piezoelectric polymers, ceramics, or composites are increasingly being investigated as powerful platforms for tissue engineering, tailored to replicate and enhance the unique electromechanical microenvironments of diverse tissues [[339], [340], [341]].
Bone is a metabolically active organ that undergoes continuous remodeling throughout its life. It is a highly specialized tissue in which osteogenic, osteoclastogenic and endothelial lineage cells each play specific roles and undergo distinct rearrangements [342]. Bone remodeling involves the exclusion of mineralized bone by osteoclasts followed by the formation of subsequent mineralization of bone matrix through the osteoblasts. Living bone exhibits several types of electrical potentials. The bone sensitivity to electromechanical stimulation is believed to be responsible for its regeneration ability. The combination of mechanical strain stimuli and endogenous electrical currents has been shown to regulate osteogenic growth [49]. Fukada and Yasuda attributed this phenomenon to piezoelectric properties of bone [251], in which generated currents change accordingly with the moisture content, maturation state, and architectural organization [304]. In theory, this piezoelectric effect could translate an environmental stimuli into a biologically recognizable signal that control both bone formation and resorptive processes [343]. Some studies have reinforced the importance of the role of fluid flow as the primary mechanism for stress generated potentials in bone, although the exact contribution of piezoelectricity remains unclear [299]. This phenomenon explains why, under compression, collagen reorganizes its dipole and display negative charges on the surface and induce charged amino acids to promote electrostatic interaction with calcium and phosphate ions. Conversely, under tension, collagen yields predominantly positive changes, that influences the streaming potential and mineralization process [344,345]. Extracellular calcium is one of the main factors and plays a crucial role in regulating bone remodeling, as its concentration directly affects both bone resorption and deposition [346].
Currently available biomaterials for bone replacement and reinforcement generally do not take into consideration the inherent piezoelectricity of bone, which may disrupt its natural electrophysiologic mechanisms. Moreover, the stress applied to these biomaterials has been associated with bone resorption, which in turn leads to implant instability and consequent fracture, and more challenging revision surgeries. Ideally, a bone tissue engineering scaffold would combine both sensing and osteoinduction properties by means of a piezoelectric scaffold, to maintain physiological levels of strain at the implant site.
Osteogenesis induced by electrical currents has been extensively documented. Early techniques were mainly based on the application of external fields, and the small electrical current generated in living tissues ultimately stimulating an effective bone formation [[347], [348], [349], [350]]. The observation of piezoelectric properties focused on bone and other biopolymers stimulated similar interest in piezoelectric properties of synthetic materials such as natural and synthetic polypeptides and fluorinated polymers with the potential of their implantation envisioning the tissue regrowth (Table 5). For example, Fukada et al. [351] electrically poled Teflon films and implanted them in rabbit's femur. The authors observed callus formation and growth at both sides of the film, resulting in the formation of rigid bone. In another study, Inoue and Ohashi used Teflon and poly-γ-methyl-L-glutamate (PMLG) films to induce bone growth in rat femurs [352] and observed induced osteogenesis. However, the newly formed bone near the Teflon film had mostly been resorbed, while near the PMLG it continued to grow. The authors attributed these differences to charge decay in Teflon film, in contrast to the sustained piezoelectric properties of PMGL.Table 5Applications of piezoelectric materials in bone tissue engineering.Table 5MaterialValidation in vitroValidation in vivoReferenceTeflon films–Induced callus formation and growth in rabbit femur[351]Teflon and PMLG films–Induced osteogenesis in rat femurs; PMLG sustained growth while Teflon resorbed[352]PVDF films–Improved osteogenesis in rat tibia from 1 to 6 weeks[353]PVDF tubes–Bone tissue grew preferentially on surface due to residual charge[354]PVDF membranesEnhanced goat marrow cells adhesion, proliferation, and differentiation into osteoblasts–[355]PVDF films (converse effect)Used to strain bone cells–[[356], [357], [358]]PVDF membranesInhibitory effects on cell proliferation and differentiation–[359,360]PVDF (osteotomy experiments)–Formed fibrous capsule; electrical stimulation increased bone mass and mineral stability[[361], [362], [363]]Poled-PVDF filmsInfluenced fibronectin adsorption; supported higher pre-osteoblasts–[364]Titanium-coated PVDF filmsInduced higher cell proliferation under static and dynamic conditions–[365]PVDF composites (CoFe2O4, Zeolite NaY, CNT, montmorillonite)Zeolite/clay promoted cell response; ferrite/silver/CNT reduced viabilityNo pro-inflammatory effects for zeolite/clay[366]PVDF-TrFE nanofibers with hydroxyapatiteEnhanced osteogenic gene expression, hMSC differentiation, calcium deposition–[367]PVDF 2D/3D structuresHigher MC3T3-E1 proliferation in 2D non-porous static and 2D porous mechano-electrical–[368]PVDF with ZIF-8 nanocrystalsInduced angiogenesis and osteogenic activityRegeneration of calvarium bone defects in rats[370]ZnO/PCL/PVDF scaffoldsInduced MC3T3-E1 differentiation; inhibited bacterial growthAccelerated regeneration in rat mandibular defects[371]ZnO in PCL/PVDF (0.5 wt%)Induced cell attachment, proliferation, calcium mineralization–[372]PHB/ZnO nanofiber-aerogelPromoted BMSC osteogenic differentiationPromoted bone regeneration in rat shin defects; accelerated by US[373]PLLA films–Significantly increased ossification on rabbit tibiae periosteum[374]PLLA rods–Significantly improved callus formation in fractured feline legs[121]PLLA filmsPromoted fibronectin adsorption; improved bone-like cell adhesion, spreading, proliferation (negatively charged)–[341,364]3D printed PLLA scaffoldsWith fibrinogen coating, supported cell adhesion and proliferation–[375]PLLA with functionalized boron nitride and pluronicImproved MC3T3-E1 adhesion, proliferation, osteogenic differentiation–[376]PLLA with polyaniline/CNT or PLLA/PVDF in hydrogelPronounced osteogenic differentiation–[377,378]BaTiO3 nanoparticles in PVDFIncreased beta phase fraction; augmented output voltage–[379,380]Ba0.9Ca0.1TiO3 in PVDFImproved bioactivity and mechanical properties–[380]PCL/BaTiO3 compositesInduced osteogenic differentiation under US–[381]PCL with BaTiO3 or BNNTIncreased tensile strength; induced osteogenic differentiation under US–[382]BaTiO3/PLAHighest d33 at 20 % BaTiO3–[178]PLLA with Ca/Mn co-doped BaTiO3 (CBMT)Augmented osteogenic activity–[383]BaTiO3/PHBHigher Young's modulus, compressive strength, d31 at 20/80–[384]HA in PHBV with BaTiO3Improved hydrophobicity, roughness, mechanicalAccelerated bone formation in rat cranial defect[385]BaTiO3/HA 3D scaffoldsHigh cytocompatibility–[386]BaTiO3/calcium phosphateImproved handleability; compressive strength inverse to BaTiO3–[387,388]BaTiO3 coated Ti6Al4V–Promoted bone formation, M2 polarization, angiogenesis in sheep[[389], [390], [391]]
Marino et al. [353] hypothesized that piezoelectrically induced surface charge on polyvinylidene fluoride (PVDF) could alter the physiological status of osteoprogenitor cells — the immediate precursors of osteoblasts located in the periosteal sheath. In their study, PVDF films were implanted in the interosseous membrane of the anterior tibia of Sprague-Dawley rats. Results showed improved osteogenesis on the piezoelectric films from 1 to 6 weeks, with an increased periosteal response during the first week. Callegari and Belangero [354] further investigated the influence of PVDF tubes on bone regeneration, finding that bone tissue grew preferentially on the surface of the tubes, likely due to residual surface charge rather than a stress generated potential difference. Rodrigues et al. [355] demonstrated that PVDF membranes enhanced goat marrow cells (GMCs) adhesion, proliferation, and differentiation into osteoblasts in vitro, with new structures consisting mainly of calcium phosphate confirming the successful differentiation of GMCs into osteoblasts. PVDF films have also been proposed as a means to directly strain bone cells by converse piezoelectric effect, whereby an applied electric field produces mechanical deformation [[356], [357], [358]]. Additional strategies have been pursued to enhance scaffold functionality. For example, PVDF films dip-coated with poly(methyl methacrylate) and Bonelike® microparticles were used in order to ensure adhesion of the osteoblasts to the device surface and ensure proper electric insulation. Results indicated that cells were sensible to low substrate deformation values in a reproducible manner, and the scaffold negatively affected cell viability and proliferation. Tabary and Hung [359,360] reported inhibitory effects on cell proliferation and differentiation when using PVDF membranes with adherent cell lines. Osteotomy experiments in sheep femur and tibia [361] resulted in the formation of a fibrous capsule. However, when subjected to electrical stimulation total bone mass and mineral stability were significantly higher in cultured substrate [362,363].
The effect of piezoelectric PVDF films on protein adsorption and added cell response was also investigated. Studies showed that both positive and negative poled-PVDF films influenced the adsorption, distribution and conformation of fibronectin (FN) on the material surfaces [364]. These authors also reported that piezoelectric films supported significantly higher number of pre-osteoblasts when compared with non-piezoelectric. Moreover, the titanium-coated-PVDF piezoelectric films – which possess increased roughness and hydrophobicity, were also shown to induce higher cell proliferation under static conditions, that was further improved under dynamic conditions [365]. The application of PVDF-based composites obtained by the addition of specific fillers, such as CoFe2O4, Zeolite NaY, carbon nanotubes, and montmorillonite has been developed and studied [366]. In vitro and in vivo evaluations revealed that zeolite and clay composites promote cell response without showing in vivo pro-inflammatory effects. On the other hand, ferrite, silver composites and carbon nanotubes showed reduced osteoblasts and fibroblasts viability, respectively.
Additional PVDF scaffold fabrication methods have been studied recently. Barbosa et al. developed PVDF-TrFE nanofibers filled with hydroxyapatite [367] which promoted β phase nucleation and improved both piezoelectric and osteoinductive properties. The nanoparticles enhanced osteogenic gene expression, human mesenchymal stem cell (hMSCS) osteogenic differentiation and calcium deposition (Fig. 3b). Silva et al. [368] used solvent casting and doctor blade to develop 2D and 3D structures with and without porosity. Results showed that higher cell proliferation of MC3T3-E1 pre-osteoblasts was achieved when using a 2D non-porous scaffolds in static culture conditions and a 2D porous scaffolds under mechano-electrical conditions (Fig. 3c). Song et al. [369] employed an in-situ chemical foaming assisted fused deposited modeling method to fabricate 3D PVDF bone-like structures with arbitrary shape design and internal pores that amplified the scaffolds piezoelectric potential (Fig. 3d). Foam fabrication was similarly used by Chen et al. [370] to fabricate PVDF scaffolds that were then combined with ZIF-8 nanocrystals. The resulting piezoelectric potential and release of Zn^2+^ from the foam strcutures induced angiogenesis and osteogenic activity in vitro and also regeneration of calvarium bone defects in rats (Fig. 3e).Fig. 3Applications of piezoelectric materials in bone tissue engineering. a) Representation of the 3-step preparation (electrostatic spinning, homogenizing and freeze-drying) of piezoelectric CS/PHB@ZnO nanofiber-aerogel scaffold and its osteogenic effect driven by ultrasound. Reproduced with permission from Ref. [373]. b) Representation of the development process of hydroxyapatite-filled osteoinductive piezoelectric PVDF-TrFE electrospun nanofibers as a strategy for bone repair in osteoporotic-related fractures. Reproduced with permission from Ref. [367]. c) development of piezoelectric 2D and 3D porous scaffolds, mimicking bone's structure and mechano-electric microenvironment, for bone tissue regeneration. Reproduced with permission from Ref. [368]. d) 3D printing of piezoelectric parts that amplify the stress–strain effect and improve the output capacity of bioinspired bone structures of PVDF. Reproduced with permission from Ref. [369]. e) The application of piezoelectric PVDF/ZIF-8 foam-based sheets promote new bone formation via coupling of angiogenesis and osteogenesis. Reproduced with permission from Ref. [370]. f) Process of surface modification of barium titanate nanoparticles, involving dispersion in water, dispersion in dopamine and polymerization, before incorporation in PVDF/composite scaffolds. Reproduced with permission from Ref. [379]. g) Fabrication process for Ca/Mn co-doped Barium titanate nanofibers composited with PLLA to strengthen the piezoelectric effect of scaffolds and improved osteogenic activity. Reproduced with permission from Ref. [383]. h) Preparation process of an artificial piezoelectric periosteum to guide bone healing under piezoelectric stimulation. Reproduced with permission from Ref. [385]. i) A bioactive composite scaffold consisting of piezoelectric barium titanate, coated on porous Ti6Al4V showed potential in promoting bone and blood vessel ingrowth in reconstructing an electromagnetic microenvironment. Reproduced with permission from Ref. [390]. j) Ti6Al4V scaffolds coated with piezoelectric barium titanate implanted in a sheep cervical en-bloc spondylectomy model promoted M2 polarization of macrophages and bone repair, through inhibition of MAPK/JNK signaling and activation of OXPHOS of macrophages. Reproduced with permission from Ref. [391].Fig. 3
Zinc, in the form of ZnO, has been incorporated into composite materials to take advantage of its piezoelectric properties and antibacterial properties which are similar to those of bone. Nanofibrous scaffolds composed of ZnO, PCL and PVDF demonstrated piezoelectric potential similar to healthy bone [371]. Moreover, these composites also inhibited bacterial growth and induced the differentiation of MC3T3-E1 cells under ultrasound (US) stimulation. In rat mandibular defect models, employing these scaffolds also accelerated regeneration. Using the same materials combination Ghaedsharafi et al. [372] recently found that incorporating 0.5 wt% ZnO nanoparticles produced the highest β phase fraction in these composites, although it reduced the tensile strength without affecting the ability of scaffolds to induce cell attachment, proliferation and calcium mineralization.
Other biodegradable polymers have also been studied. PHB is superior to PVDF in terms of nontoxicity and biodegradability though, the former has lower piezoelectric coefficients than the latter. In order to enhance the potential for PHB to generate electrical fields, Chen et al. [373] prepared a 3D nanofiber-aerogel scaffold fabricated with PHB and ZnO nanofibers. The addition of 2 % ZnO resulted in scaffolds with optimal piezoelectric and mechanical properties, promoting in vitro differentiation of osteogenic bone marrow stromal cells (BMSC). Additionally, the implantation of the nanofiber-aerogels in rat shin bone defects promoted bone regeneration that was accelerated by US stimulation (Fig. 3a).
Poly(ʟ-lactide acid) (PLLA) films were developed in an attempt to improve scaffold biodegradability. Using this strategy, PLLA films achieved significantly increased ossification on the periosteum of rabbit tibiae [374] and Ikada et al. [121] reported significantly improved callus formation in fractured feline legs fixed using PLLA rods suggesting that the piezoelectric current generated by mechanical strains that accompanies leg movement of the cats was a key factor for bone formation. Moreover, piezoelectric PLLA films were shown to promote increased fibronectin adsorption over paraelectric films. These films have also improved bone-like cell adhesion, spreading and proliferation, particularly when the films are negatively charged [341,364].
Modern technologies, such as 3D printing, have been recently employed to fabricate PLLA scaffolds. For instance, Karanth et al. [375] used fused deposition modelling technology to produce scaffolds with 73 % porosity, an average pore size with of 450 μm and able to generate an electric potential of 25 mV. Fibrinogen coating on the scaffolds was necessary to support cell adhesion and proliferation. Ramasamy et al. [376] demonstrated that re-enforcement of PLLA matrices with functionalized boron nitride and pluronic results in composites with enhanced mechanical and piezoelectric performance that improve the adhesion, proliferation, and osteogenic differentiation of MC3T3-E1 cells. Additionally the group of Hajrezaei and Samghabadi [377,378] used electrospinning technique to develop PLLA scaffolds blended with other materials. They prepared two groups of the first included PLLA, polyaniline and carbon nanotubes, while the second group embedded PLLA and PVDF mats into a hydrogel made of oxidized alginate, gelatin and poly-pyrrole-graft-gelatin. In both cases, the composites presented high piezoelectric capability and biocompatibility. Cells cultured on both composites showed pronounced osteogenic differentiation.
Bone's native composition, mostly collagen and hydroxyapatite, has in recent years spurred the use of piezoelectric composites, using bioactive ceramics in bone tissue engineering. Among bio ceramics BaTiO3 is one that is frequently employed although its poor processability and reduced toughness presents challenges. One strategy of overcoming these limitations is combining piezoelectric ceramics and polymers. For example, using BaTiO3 nanoparticles to enhance piezoelectric features of other materials has been a strategy adopted by multiple research groups. Shuai and Ahmadi [379,380] dispersed BaTiO3 nanoparticles functionalized with polydopamine in PVDF and achieved an increase from 46 % to 59 % in β phase fraction, while also augmenting the scaffolds output voltage of by 365 % (Fig. 3f). Ahmadi et al. [380] demonstrated that uniformly dispersed of Ba0.9Ca0.1TiO3 in PVDF resulted in improved bioactivity and mechanical properties of membranes.
Composite formulations combining polycaprolactone (PCL) with BaTiO3 have similarly been developed. Skider et al. [381] produced PCL/BaTiO3 composites in the shape of uniform diameter filaments as feed stock for fused filament fabrication based 3D printing. Çobandede et al. [382] employed electrospinning to embed BaTiO3 or boron nitride nanotubes (BNNT) into PCL obtaining fibrous structures that increased the tensile strength of PCL while decreasing its hydrophobicity. In both studies, scaffolds could generate electric fields in response to external US stimulation, thereby inducing osteogenic cell differentiation.
Dai and colleagues prepared piezoelectric composite membranes composed of BaTiO3 and Polylactic acid (PLA), and obtained the highest piezoelectric coefficient (d33) with 20 % concentration of BaTiO3 [178]. Zheng et al. [383] combined PLA stereoisomer, PLLA, with Ca/Mn co-doped BaTiO3 (CBMT) nanofibers which augmented osteogenic activity (Fig. 3g). However, the properties of the scaffolds declined with increased CMBT content.
Strangis et al. [384] used BaTiO3 nanoparticles as a filler for a PHB matrix. Among different w/w% compositions tested, 20/80 BaTiO3/PHB produced higher Young's modulus, compressive strength and piezoelectric coefficient d31. Liu et al. [385] further incorporated hydroxyapatite (HA) into poly(3-hydroxybutyric acid-co-hydrovaleric acid (PHBV) polymers reinforced with using a one-step spin-coating method to fabricate a piezoelectric periosteum (Fig. 3h). With this strategy, improved hydrophobicity, roughness and mechanical performance was achieved. In vivo testing showed that the biomimetic periosteum accelerated bone formation in a rat cranial defect model. Other researchers [[386], [387], [388]] have explored the combination of BaTiO3 and HA or calcium phosphate in the development of scaffolds. 3D scaffolds composed of a homogeneous mix of BaTiO3 and HA, fabricated by binder jetting process, present piezoelectric properties in the range of dry bone. Despite the high cytocompatibility the scaffolds fabricated they were found to be brittle due to the high microporosity [386]. Addition of BaTiO3 to calcium phosphate resulted in improved scaffold handleability [387] but compressive strength was affected in an inverse relationship to BaTiO3 concentration. In turn, the increase in calcium phosphate phase results in an increase of the dielectric constant [387,388].
Titanium scaffolds have commonly been used in the repair of bone defects due to their ability to provide mechanical strength which contributes to maintaining bone defect stability, while its porous structure provides support for tissue growth. However, because titanium is a bioinert material, its ability to bind to bone might be compromised over time, potentially leading to long term implantation failures [389]. To address this issue, surface modification of titanium alloys such as by coating porous Ti6Al4V with BaTiO3 have been investigated. The coating modified scaffolds were able to create an electrically favorable microenvironment and provided mechanical properties suitable for bone regeneration. In sheep models, studies revealed that BaTiO3 coated scaffolds were more successful in promoting bone formation and enhanced M2 macrophage polarization and angiogenesis were observed (Fig. 3i and j) [[389], [390], [391]].
The precise mechanism by which bone tissue recognizes the piezoelectric stimuli and transforms it into biological activity remains unclear. Fukada [392] provides a possible explanation of the osteogenic power of these materials. He hypothesized that due to the piezoelectric polarization induced by muscular activity or subtle bone movements, charges are generated. These charges can produce a flow of ionic currents in the tissue fluid surrounding the implant, which is electrically conductive. It is postulated that these ionic currents might represent a stimulus for cell migration or even differentiation of fibroblasts into osteoblasts, providing a fast proliferation of bone cells. Alternatively, mechanical deformation at physiological levels may create localized hydrostatic pressure gradients in the interstitial fluid of the caniculi in the bone. These gradients would be further captured by the cellular membrane of the osteocyte, resulting in increased metabolic activity, suggesting that the transduction process from the mechanical to the biological stimuli is regulated by these cells [354].
Cartilage is a resilient and avascular tissue that provides low-friction surfaces for joint movement and absorbs mechanical loads. However, due to its limited regenerative capacity, cartilage injuries often lead to progressive degeneration and osteoarthritis (OA), posing significant clinical challenges [393,394]. Current treatments, such as analgesics, anti-inflammatory drugs, microfracture, or autologous chondrocyte implantation, offer symptomatic relief but fail to restore native hyaline cartilage structure and function. Piezoelectric materials have emerged as promising scaffolds in cartilage TE by generating localized electrical cues that mimic the endogenous electromechanical microenvironment of cartilage, which naturally produces piezoelectric charges under deformation. These materials can promote chondrocyte proliferation, ECM synthesis, and stem cell differentiation toward chondrogenic lineages, potentially addressing the limitations of traditional approaches. Recent studies have explored various piezoelectric strategies, including US-activated hydrogels, motion-induced scaffolds, injectable nanofibers, piezo-triboelectric generators, and piezoionic systems, to enhance cartilage regeneration [[395], [396], [397]].
Several studies have focused on injectable piezoelectric hydrogels activated by US stimulation, leveraging the deep tissue penetration and non-invasive nature of US to generate electrical signals in situ (Table 6). These hydrogels often incorporate piezoelectric nanoparticles or polymers to convert mechanical US waves into bioelectric cues that promote chondrogenesis and modulate inflammation. One recent strategy involved embedding piezoelectric nanoparticles in biocompatible hydrogel matrices. For instance, a piezoelectric hydrogel (Hyd6) was fabricated using polyvinyl alcohol (PVA), polyacrylic acid (PAA), chitosan, and KNN nanoparticles through dynamic hydrogen bonding and covalent cross-linking [398]. This hydrogel exhibited injectability, self-healing, adhesion, and electromechanical conversion properties. In vitro experiments with rabbit bone marrow mesenchymal stem cells (rBMSCs) demonstrated that Hyd6 under US stimulation enhanced cell recruitment, Ca^2+^ influx, and activation of the CaM/CaN/NFATc1 pathway, leading to upregulated expression of chondrogenic genes (SOX9, COL2A1, ACAN). In vivo, in a rabbit chondral defect model, Hyd6 plus US treatment after 8 weeks showed improved cartilage integration, higher collagen type II deposition, and better mechanical properties compared to controls, highlighting its potential as a self-powered therapy for cartilage repair.Table 6Applications of piezoelectric materials in cartilage tissue engineering.Table 6MaterialValidation in vitroValidation in vitroReferencePVA/PAA/chitosan/KNN hydrogel (Hyd6)Enhanced rBMSC recruitment, Ca2+ influx, chondrogenic gene expressionImproved cartilage integration, collagen II deposition in rabbit defects[398]HA with Ga/PDA/BT (PBT-Gel)Enhanced chondrogenesis via TGF-β/Smad; anti-inflammatory M1 to M2Reduced synovitis, cartilage degradation in rat OA[399]BaTiO3/GO in polysaccharide matrixUpregulated chondrogenic genes; downregulated fibrotic/catabolic; anti-inflammatory–[400]MOF(Hf)-Pt nanozymeScavenged ROS, restored mitochondrial in rat chondrocytesReduced cytokines, OARSI scores in rat OA[401]SrSiO3/PLLAEnhanced chondrogenic/osteogenic differentiation via P2RX1-Ca2+Superior integration, higher COL2/OCN in rat defects[402]GeSe nanosheets (GSN)–Reduced cytokines, alleviated synovitis/cartilage erosion in rat RA[403]PLLA nanofiber scaffold–Promoted rBMSC migration, chondrogenesis; hyaline regeneration in rabbit defects with exercise[404]PLLA with DANCR-EVsEnhanced EV uptake, organized collagen IIFormed cartilage lacunae in rat defects[405]PLLA short nanofibers with GO/PMPCImproved mitochondrial function in OA chondrocytesReduced friction, inflammation in rat OA[406]PHBV/PLLA composite–Enhanced remodeling via Ca2+/TGF-β; superior hyaline formation in rabbit defects[407]PVDF/ZnO/PCL (PZP)–Promoted proliferation, ECM synthesis; strong regeneration in rat defects with exercise[408]SA/PHBV/Fe3O4 hydrogel–Upregulated chondrogenic genes; hyaline regeneration in rabbit defects[409]PLLA/BT with collagen/FGF-18Promoted chondrocyte proliferation, ECM synthesisEnhanced regeneration in rabbit defects with exercise[410]PLLA nanofibers in collagen hydrogel–Induced TGF-β1, chondrogenesis; improved hyaline structure in rabbit OA[411]Gelatin/PVA with BN (PTEG)Enhanced BMSC migration, chondrogenesisReduced osteophytes in rat OA with exercise[412]GelMA/polyrotaxane/Ppy/BTO microspheresPromoted chondrogenic differentiation, M2 polarizationAlleviated cartilage damage in rat OA[413]Gel-PC/dECM/FF/PEDOT gelPromoted BMSC biphasic differentiationEnhanced repair via electrical gradients in Parma pig defects[414]PCL/G with PVDF/T-BT–Recruited stem cells, maintained ECM, promoted regeneration in rat IVD[415]
Similar approaches used BaTiO3 nanoparticles in hyaluronic acid (HA)-based hydrogels. A nanocomposite hydrogel (PBT-Gel) was developed by coordinating gallium (Ga^3+^) ions with HA carboxyl groups and incorporating polydopamine (PDA)-modified BT nanoparticles [399]. The hydrogel showed injectability, biodegradability, and antioxidant properties. In vitro tests with rBMSCs and RAW264.7 macrophages under US revealed enhanced chondrogenesis via TGF-β/Smad signaling and anti-inflammatory effects through macrophage repolarization from M1 to M2 phenotype (Fig. 4d). In a rat OA model, PBT-Gel plus US reduced synovitis and cartilage degradation after 4 weeks, as shown by lower OARSI scores. Another US-responsive hydrogel incorporated BaTiO3 nanoparticles and GO nanoflakes in a polysaccharide matrix [400]. Stimulated with US it boosted chondrogenic differentiation of human adipose-derived stromal cells in vitro, upregulating SOX9, COL2A1, and ACAN while downregulating fibrotic and catabolic markers. Proteomic analysis showed influences on cytoskeleton organization, ECM remodeling, and metabolic processes, with a notable anti-inflammatory effect in an IL-1β milieu.Fig. 4Applications of piezoelectric materials in cartilage tissue engineering. a) Ultrasound-responsive lubricating electroactive short fibers upregulate Sirt1 gene expression and restore mitochondrial quality control mechanisms resulting in repair of the hydration lubrication layer of cartilage and improving osteoarthritis. Reproduced with permission from Ref. [406]. b) Diagram of the microstructure of normal articular cartilage used as model for bionic PLLA scaffolds for cartilage regeneration. Reproduced with permission from Ref. [405]. c) Diagram of the fabrication of 0.7PHBV porous scaffolds, made of 70 % PHBV and 30 % PLLA, wt%), ensuring low Young's modulus and additional piezoelectricity. Reproduced with permission from Ref. [407]. d) Application of a PBT hydrogel with ultrasound resulting in immunomodulatory macrophage repolarization, reduced inflammation and cartilage regeneration. Reproduced with permission from Ref. [399]. e) A piezoelectric pPLLA/SrSiO3 composite scaffold sustainably releases bioactive Sr^2+^ and SiO3^2−^ ions, that in combination with LIPUS stimulation, promotes osteogenic differentiation of rBMSCs and enhances cartilage regeneration and subchondral bone reconstruction. Reproduced with permission from Ref. [402].Fig. 4
Piezoelectric metal-organic frameworks (MOFs) have also been explored. A US-triggered piezocatalytic nanozyme (MOF(Hf)-Pt) was synthesized by anchoring platinum (Pt) nanoparticles on hafnium-based MOFs [401]. Under US, it exhibited enhanced superoxide dismutase and catalase activities, scavenging reactive oxygen species and restoring mitochondrial function in rat chondrocytes. In a rat OA model, MOF(Hf)-Pt plus US reduced inflammatory cytokines (TNF-α, IL-1β) and OARSI scores by 76.90 % after 4 weeks, promoting cartilage matrix synthesis. Liu et al. [402] combined Strontium silicate (SrSiO3) bioceramics with PLLA to form composite scaffolds. Fabricated by electrospinning, the PLLA/SrSiO3 scaffold released Sr^2+^ and SiO3^2−^ ions and generated piezoelectric charges under low-intensity pulsed US. In vitro, it enhanced chondrogenic differentiation of rat chondrocytes and osteogenic differentiation of rBMSCs via P2RX1-mediated Ca^2+^ signaling. In a rat osteochondral defect model, pPLLA/5 %SrSiO3 plus LIPUS after 8 weeks showed superior integration, with higher COL2 and OCN expression (Fig. 4e).
Sonocatalytic nanosheets represent another US-activated strategy. Zeng et al. [403] followed this approach by performing intra-articular injection of germanium selenide (GeSe) nanosheets (GSN) with piezoelectric properties. Under US, GSN catalyzed H2 production and lactic acid consumption, suppressing Th1 polarization and osteoclastogenesis in a rat RA model induced by complete Freund's adjuvant. After 21 days, it reduced TNF-α, IFN-γ, and RANKL levels, alleviating synovitis and cartilage erosion. Piezoelectric scaffolds activated by joint motion or exercise harness physiological mechanical loads to generate electrical cues, eliminating the need for external stimulation devices. Liu et al. [404] implanted a biodegradable PLLA nanofiber scaffold in rabbit osteochondral defects. Under treadmill exercise it generated piezoelectric charges, promoting rBMSC migration, Ca^2+^ influx, and chondrogenesis via endogenous TGF-β1 induction. After 1–2 months, the scaffold plus exercise group showed hyaline cartilage regeneration with abundant type II collagen. In a strategy combining piezoelectric stimulation with extracellular vesicles (EVs), Lai et al. [405] tested a PLLA scaffold loaded with DANCR-overexpressing synovium-derived MSC-EVs. In vitro, piezoelectric signals enhanced EV uptake by chondrocytes, organizing type II collagen along PLLA fibers. In a rat articular cartilage defect model, PLLA/EVs after 6 weeks formed cartilage lacunae, facilitating chondrocyte residence and regeneration (Fig. 4b). Injectable PLLA short nanofibers modified with GO and poly[2-(methacryloyloxy)ethyl phosphorylcholine] (PMPC) were developed by Ding et al. [406] to form electroactive fibers. Under US, they generated 1.31 V, upregulating Sirt1 in OA chondrocytes and improving mitochondrial function. In a rat OA model, the fibers reduced friction coefficient to 0.235 and alleviated inflammation after 4 weeks (Fig. 4a). 3D printing of scaffolds was employed in recent works by two groups. A biodegradable piezoelectric polymer composite (0.7PHBV, 70 % PHBV/30 % PLLA) was 3D-printed into porous scaffolds [407], resulting in enhanced cartilage remodeling via Ca^2+^ and TGF-β signaling in rabbit knee defects under natural motion, and showing superior hyaline cartilage formation after 8 weeks (Fig. 4c). Employing 3D printing and rolling, Lin et al. [408] fabricated Hierarchically structured PVDF/ZnO/PCL (PZP). In rat cartilage defects, PZP plus exercise generated ∼5 V, promoting chondrocyte proliferation and ECM synthesis via TGF-β pathway. After 1–2 months, it achieved strong cartilage regeneration. A magnetoelectric hydrogel composed of SA, PHBV and Fe3O4 was assembled to mimic cartilage structure [409]. In rabbit defects, motion-driven magnetoelectric coupling produced enhanced electromagnetic stimuli, upregulating TGF-β, SOX9, ACAN and COL2A1 and promoting hyaline cartilage regeneration after 8 weeks.
Injectable systems allow minimally invasive delivery and conform to irregular defects. Following this strategy, Xie et al. [410] integrated PLLA/BT nanofibers with collagen and FGF-18 in a 3D composite scaffold. In vitro assays authors showed that it promoted rabbit chondrocyte proliferation and ECM synthesis. In rabbit defects, the scaffold plus exercise enhanced cartilage regeneration via piezoelectricity and biofactor synergy after 8 weeks. Additionally, Vinikoor et al. [411] injected PLLA nanofibers in collagen hydrogel into rabbit OA defects. Under US it induced TGF-β1 secretion and chondrogenesis, improving hyaline cartilage structure after 2 months.
Hybrid mechanisms have been tested combining piezoelectric and triboelectric effects for higher charge output. Recently, Zhang et al. [412] developed a biodegradable piezo-triboelectric generator (PTEG) using gelatin/PVA with boron nitride nanosheets. Generating 242 μC/m^2^ in vitro and 84 μC/m^2^ in vivo under joint load, it enhanced BMSC migration and chondrogenesis. In rat OA models, PTEG plus exercise reduced osteophytes by 50 % after 8 weeks. Wang et al. [413] developed slide-ring structured hydrogel microspheres using GelMA, polyrotaxane, Ppy and BTO) that resulted in reduced energy dissipation to 43 %. After testing in a rat OA, authors observed that the hydrogel promoted chondrogenic differentiation and M2 macrophage polarization via cAMP signaling, alleviating cartilage damage. Another piezoelectric gel was developed by Liu et al. [414]. The gel was constituted of Gel-PC, dECM, FF and PEDOT. Authors showed that it promoted BMSC biphasic differentiation, and that in Parma pig osteochondral defects it enhanced repair via electrical gradients after 12 weeks. For intervertebral disc (IVD) repair, Wang et al. [415] followed an approach based on a sandwich-structured PCL/G with PVDF/T-BT scaffold that generated dual-electroactive cues. Testing in rat IVD defects post-discectomy showed that it recruited stem cells and maintained ECM, promoting regeneration after 8 weeks.
In summary, piezoelectric scaffolds offer a versatile platform for cartilage TE, with US or motion activation generating bioelectric cues that enhance chondrogenesis, reduce inflammation, and promote ECM remodeling. Hydrogels and injectable systems dominate for minimally invasive delivery, while composites with biofactors or hybrid mechanisms were shown to provide amplified effects. In vivo studies in rat and rabbit models consistently show improved hyaline cartilage regeneration, underscoring clinical potential. Future work should focus on long-term biocompatibility, scalable fabrication, and human trials to translate these innovations.
Nerve regeneration is a complex biological phenomenon. Inputs from muscles, involuntary organs, and senses travel through the nerves of the peripheral nervous system (PNS) to the central nervous system (CNS) where they are processed and interpreted. After damage, the nerves in the PNS are often capable of spontaneously healing, depending on the size of the defect. However, when severe lesions occur, this system does not spontaneously heal due to incomplete axonal regeneration leading to permanent patient disability [416]. Because neuronal cells exhibit electrical activity, the transfer of bioelectrical signals is essential for proper nerve function and recovery. For this reason, an ideal peripheral nerve scaffold should not only mimic the fine structure of natural nerves but also replicate their electrical properties. Piezoelectric materials, which can transform mechanical stress into electrical signals, have the ability to stimulate intracellular signaling pathways vital for cell activity and function. This makes them promising candidates for promoting nerve tissue regeneration. Consequently, in recent years several research groups have put efforts in developing novel piezoelectric scaffolds that are capable facilitating nerve repair and axonal growth (Table 7) [6,417].Table 7Applications of piezoelectric materials in neural tissue engineering.Table 7MaterialValidation in vitroValidation in vitroReferencePVDF-TrFE tubesEnhanced Nb2a differentiation, neurite outgrowthIncreased efficiency of nerve regeneration; More myelinated axons[419,420]PVDF channels–Supported neuronal morphology, reduced cell death in rat PNS[421]PVDF filmsIncreased neuronal density, survival, neurite length/branching in rat spinal cord–[422]PVDF nanostripesEnhanced adhesion, migration, neuronal markers in rbMSCs–[423]PVDF conduits–Greater myelination of axons in sciatic nerve[424]PVDF-TrFE aligned matsEnhanced neurite extension in dorsal root ganglion–[425,426]Poled PVDF tube with collagen/NGFAssisted fiber regeneration, reduced cell death–[428]PVDF with silk/MXene or PEDOTInduced growth, proliferation, myelination in Schwan cellsRegeneration and recovery in rat sciatic[429,430]Poled PLGAEnhanced neurite outgrowthMore myelinated axons, faster conduction in rat sciatic[431]ZnO in polyurethaneReduced astrocyte adhesion, proliferation–[432]ZnO/PGA in chitosan-gelatin with PCL/PVDFInduced Nestin/MAP2 in PC12; antibacterial–[433]ZnO/PCL nanofibers–Improved sciatic repair, functional recovery in rats[434]PCL with boron nitride nanosheetsModulated Schwan viability, neuron regrowthRepaired sciatic defect in rats[435]PVDF-GOImproved PC12 attachment, spreading, proliferation–[436]BaTiO3/dopamine/PVDF–Stimulated nerve growth, restored function in rat sciatic[437]BaTiO3 in PVDF-TrFE with PNIPAAmIncreased neuronal markers, neurite length, myelination in PC12/SchwanNerve growth, restore function in rat sciatic[438]KNN nanowires with PLLA/PHBV/PLA/PCL–Nerve growth, restore function in rat sciatic[439]PCL/PVDF with nanotopographyIncreased markers, neurite in PC12Growth, function in rat sciatic[440]CoFe2O4 in PVDF-TrFE helicalPromoted PC12 differentiation–[441]Fe3O4/PVDFPromoted neuronal orientation, proliferation, migration, axon growthEnhanced sciatic repair, functional recovery in rats[442]PVDF-TrFE fibrousEnhanced differentiation, neurite; maintained nestin+ in hNSCs/NPCs–[443,444]PVDF-BaTiO3 nanoclusters–Dopaminergic neuron regeneration, behavioral improvement in zebrafish/mouse Parkinson's[445]
PVDF and its copolymer polyvinylidene fluoride-trifluoroethylene (PVDF-TrFE) are semi-crystalline polymers where the former becomes piezoelectric if it is in an all-trans oriented dipole configuration, displaying a degree of piezoelectric activity than exceeds any biological material [418]. One study investigating the piezoelectric properties of PVDF measured using a 1200 Hz oscillating voltage with a of 2.5 mV peak [419], demonstrated that neural cell (Nb2a) differentiation and neurite outgrowth were enhanced compared to paraelectric controls. In both in vitro and in vivo PVDF showed an increased efficiency of nerve regeneration [419,420]. 10.13039/100014337Furthermore, PVDF based channels have shown the ability to support neuronal morphology and reduced cell death when used to repair in rat models of PNS injury [421].
Royo-Gascon et al. [422] explored the use of PVDF films both in their piezoelectric and paraelectric state as in vitro scaffolds for growing rat spinal cord neurons. They found that piezoelectric stimulation dramatically increased neuronal density, improved cell survival and attachment (likely due to greater adsorption of matrix proteins). Additionally, improved growth in terms of both neurite length, fundamental to proper neuronal development and branching complexity, that is essential for establishing connections. Complementary work by Zhang et al. [423] demonstrated that the incorporation of nanotopographic features in PVDF nanostripes array structures had a significant impact in enhancing piezoelectric potential and resulting in improved cell adhesion and migration and induced the expression of neuronal markers in rabbit bone marrow stromal cells (rbMSCs).
In studies focused on regenerating peripheral nerves, Aebischer et al. [424] proposed the use of PVDF conduits to target sciatic nerve regeneration, and reported significantly greater myelination of axons compared to unpoled-PVDF guidance channels. They also investigating PVDF-TrFE, which does not require mechanical stretching to achieve a dipole-containing crystal structure. PVDF-TrFE tubes promoted regenerated nerves with significantly more myelinated axons. Enhanced neurite outgrowth in these system appears to be due to charge generation and independent of the chemical composition [419,420]. Additionally, dorsal root ganglion neurons seeded on PVDF-TrFE electrospun aligned mats showed enhanced neurite extension [425,426].
Nerve guidance conduits have also been further enhanced in combination with relevant biochemical cues, such as nerve growth factor (NGF), brain-derived nerve growth factor, laminin, collagen, FN, and other ECM molecules in order to mimic a neurite regenerative environment and correct nerve defects [427,428]. For instance, Delaviz et al. [428] prepared a poled PVDF tube filled with collagen gel – leveraging collagen's protective mechanical properties along with NGF to use in vitro in an attempt to develop a nerve autograft substitute for repairing a 10 mm sciatic nerve defect. Although this approach was not the most optimal, authors demonstrated assisted forward regeneration of the fibers conferring a protective sheath with a reduced cell death using these autograph substitute and its potential to substitute nerve regeneration by autologous graft [428].
PVDF has also been combined with other materials for the development of nerve guidance conduits. Zhang and Ma groups [429,430] integrated silk fibroin and MXene or PEDOT in electroactive scaffolds. In both cases authors were able to induce the growth, proliferation and myelination of Schwan cells cultured on these electroactive scaffolds. Additionally, conduits promoted the regeneration and recovery of function of rat sciatic nerves Fig. 5b and c).Fig. 5Applications of piezoelectric materials in neural tissue engineering. a) aligned piezoelectric nanofibers derived hydrogel barium titanate piezoelectric nanoparticles doped PVDF-TrFE nanofibers with aligned nanofiber orientation were used to fabricate an hydrogel with integrated bioactive drug. The nanofibers are able to induce neuronal oriented extension, and generate wireless electrical stimulation that promote neurite growth and cell differentiation under mechanical stimulation of ultrasound excitation. Reproduced with permission from Ref. [438]. b) A silk fibroin, PVDF-HFP, MXene composite scaffold responds to external ultrasounds promoting Schwan cell growth and proliferation. The nerve scaffold enhanced axon elongation and favorable recovery of motor and sensory function. Reproduced with permission from Ref. [429]. c) Construction of a multi-channel silk fibroin cryogel scaffold polymerized with PEDOT and wrapped with PVDF/PCL film. The nerve guidance conduit shows potential for to improve sciatic nerve function. Reproduced with permission from Ref. [430]. d) Nanofiber tubes were fabricated by electrospinning PCL with zinc oxide and wrapped with PZNF. Implantation for nerve bridging induced Schwan cell recruitment and axon regeneration and connection. Reproduced with permission from Ref. [434]. e) Representation of boron nitride nanosheets functionalized polycaprolactone channel scaffold fabricated by 3D layer-by-layer droplet spraying method. mechanical stimulation of the piezoelectric scaffold improves muscle reinnervation and locomotor recovery. Reproduced with permission from Ref. [435]. f) Development of porous PVDF-GO nanocomposite scaffolds fabricated using NIPS process by immersion precipitation followed by freeze drying. Reproduced with permission from Ref. [436]. g) An ultrasound responsive nanogenerator was developed on basis of piezoelectric composite thin films containing 0.5Ba(Zr0.2Ti0.8)O3-0.5(Ba0.7Ca0.3)TiO3 nanowires and PVDF. The implantable piezoelectric thin film nanogenerator are actuated by remote ultrasound excitation with adjustable input power and waveform, exerting direct electrical neurostimulation in sciatic nerves of rats. Reproduced with permission from Ref. [437]. h) A piezoelectric nanogenerator composed of KNN, PLLA and PHBV with encapsulation layers of PLA or PCL and magnesium and molybdenum wires is externally activated by ultrasound and enhances sciatic nerve regeneration. Reproduced with permission from Ref. [439]. i) Electrospun aligned piezoelectric nanotracts composed of PCL and PVDF combines ultrasound-activated modality and unidirectional nanotopography-based modalities that promote sciatic nerve regeneration and restoration of motor function. Reproduced with permission from Ref. [440]. j) A pomegranate-inspired barium titanate piezoelectric nanogenerator converts ultrasound energy to electrical pulses triggering membrane depolarization and calcium influx in nerve cells, facilitating neuronal recovery in a Parkinson's disease model. Reproduced with permission from Ref. [445].Fig. 5
Bryan et al. [431] tested poly(lactic-co-glycolic acid) (PLGA) for PNS tissue regeneration by rendering it piezoelectric through a corona poling process, enabling it to generate an average voltage of 2.5 mV. PLGA films prepared under two different poling conditions were used to evaluate neurite outgrowth in vitro and nerve regeneration in vivo. Results showed that regeneration on PLGA that was poled at 24 kV for 45 min led to significantly more myelinated axons and faster conduction velocity than constructs that were poled at 20 kV for 10 min. Similar regenerative benefits were observed in scaffolds bridging a 1 cm sciatic nerve gap.
In recent years, nanomaterials have emerged as promising candidates for a variety of TE applications. Seil and Webster [432] embedded piezoelectric zinc oxide (ZnO) nanoparticles within polyurethane scaffolds to study the in vitro activity of astroglial cells on these composites. The results obtained indicate reduced astrocyte adhesion and proliferation, which is an important feature as it minimizes the glial scar formation and reduced neuronal immune response, indicating its potential as a NGC. Javidi et al. [433] further combined ZnO nanoparticles and poly(γ-glutamic acid) (PAG) in a chitosan-gelatin hydrogel used as the filling for a co-electrospinning developed nanocomposite constituted of PCL/PVDF, gelatin and polyaniline/graphene. The incorporation of PGA contributed to increased electrical conductivity but decreased biodegradability. The resulting conduit presented antibacterial activity and importantly induced the expression of neural markers Nestin and MAP2 in PC12 cells [433]. Moreover, ZnO/PCL nanofibers have been developed to serve as a piezoelectric nerve conduit. The electrical stimulation that was generated by this conduit was able to induced an increase in the number of myelin sheets and recruitment of Schwan cells (Fig. 5d). It also resulted in an improved sciatic nerve repair and faster functional recovery in rats than when using PCL alone [434]. In another study, PCL has also been employed to functionalize boron nitride nanosheets with the aim to form a piezoelectric scaffold that modulated Schwan cells viability and induced neuron regrowth while effectively repairing a sciatic nerve defect in rats (Fig. 5e) [435].
The incorporation of carbon nanostructures, such as GO in polymer matrices results in the creation of numerous surface charges that can influence nerve response in terms of neurite number and length. Abzan et al. [436] developed nanocomposite scaffolds based on PVDF-GO with the aim of exploring the optimal GO content. The incorporation of 0.5 wt% GO in the constructs resulted in improved piezoelectricity and reduced electrical conductance. When tested on PC12 cells, these scaffolds showed improved cell attachment, spreading and proliferation (Fig. 5f).
Several research groups have recently devised piezoelectric nanogenerators activated by an external energy source such as US. This approach offers advantages like deep tissue penetration, high spatial resolution, and the ability to precisely modulate the electrical-stimulation parameters. The groups of Chen and Xu [437,438] used BaTiO3 as the basis for their proposed composites. Chen et al. [437] combined BaTiO3 with a dopamine coating and a PVDF matrix to enhance piezoelectric properties (Fig. 5g), while Xu et al. [438] opted to combine BaTiO3 nanoparticles into aligned PVDF-TrFE incorporating a poly(N-isopropylacrylamide) to form an outer layer that added drug encapsulation capability. Wu et al. [439] developed a piezoelectric scaffold composed of KNN nanowires, combined with PLLA, PHBV, PLA or PCL films, and metallic additives such as magnesium and molybdenum (Fig. 5h). Pi et al. [440] created a biodegradable piezoelectric scaffold by integrating PCL and PVDF with nanotopographic features for nerve guidance. In vitro studies of Pi and Xu showed that these US responsive scaffolds have the potential for nerve regeneration therapy, as they increased neuronal marker expression in PC12 cells, increased neurite length and increased myelination in Schwan cells (Fig. 5a,i**)** [438,440]. These authors tested their systems by implanting in sciatic nerve defect of rat where the ability to stimulate nerve growth and to restore nerve motor function was observed [[437], [438], [439], [440]]. An alternative form of external electrical stimulation was developed by Chen and collaborators by dispersing CoFe2O4 magnetic nanoparticles in a piezoelectric matrix of PVDF-TrFE matrix resulting in helical magnetic microswimmers capable of transporting neuron-like cells on their surface. Additionally, acoustic wave induced piezoelectricity promoted cell differentiation of PC12 cells [441]. Another magnetically activated scaffold was elegantly devised by Wu et al. [442]. Authors employed electrospinning to develop aligned piezoelectric Fe3O4/PVDF scaffolds with good fiber orientation and controllable output signals. In vitro testing showed that the scaffolds promote neuronal cell orientation, proliferation and migration. Magnetic field also induced axonal growth and extension and increased expression of axon guidance markers. In vivo, the scaffolds enhanced sciatic nerve repair and functional recovery without inflammatory reaction.
In contrast to PNS, the lack of regeneration in the CNS derives from a complex series of events that follow injury. These include the formation of glial scar tissue that can prevent spontaneous neural regeneration by acting as physical and chemical barrier, leading to deficits of sensory and motor functions [340]. TE could offer much needed solutions to regenerate or replace damaged CNS tissue using suitable piezoelectric materials. Piezoelectric electrospun fibrous scaffold made of PVDF-TrFE fabricated at the micron- and nanoscale has been evaluated in vitro using human neural stem and neural progenitor cells (hNSCs/NPCs). Results showed enhanced cell differentiation and higher neurite extensions, while also maintaining a population of nestin + neuronal precursor cells [425,443,444]. Furthermore, in a zebrafish and mouse models of Parkinson's disease, Chen et al. [445] reported that pomegranate like PVDF- BaTiO3 nanoclusters were able to activate calcium channels and induce membrane depolarization in neurons. Moreover, zebrafish and mice presented behavioral patterns conducive of dopaminergic neuron regeneration after treatment with the nanogenerators (Fig. 5j).
The benefits of piezoelectric scaffolds for neural tissue engineering extend beyond enhancing neural cell function – they can also reduce glial cell adhesion and proliferation, thereby mitigating glial scar formation. These advances collectively underscore the potential piezoelectric material and scaffolds to not only promote nerve generation in PNS and CNS but also to provide a controllable microenvironment that supports cellular growth, differential and functional recovery.
Skeletal muscles are composed of bundles of highly oriented and dense muscle fibers-each being a multinucleated cell derived from the fusion of myoblasts [446,447]. Contractions of skeletal muscles generate the force necessary for movement, mainly of skeletal components thereby maintaining skeletal ability. They adapt to a wide variety of functional demands, and contribute to body temperature regulation. The creation of new skeletal muscle through TE approaches represents a promising alternative to muscle transfer. The process of skeletal muscle TE relies on the regenerative properties of satellite cells which proliferate and differentiate to repair and replace damaged muscle tissue [[448], [449], [450]]. Timely innervation and vascularization of TE muscle tissue are also critical in the recovery of function [[451], [452], [453]].
Biomaterials play a crucial role in supporting myogenesis because in vivo myogenic stem cells reside and differentiate within a 3D environment. In recent years, it has been described that chemical, mechanical, and topographical cues strongly affect muscle cell adhesion, proliferation, and differentiation in mature contractile structures [360,454,455]. Numerous attempts have been made to induce the fusion of myoblasts to myotubes in vitro, imitating the in vivo conditions during natural myogenesis. Critical challenges include an understanding of the effects of mechanical and electrical stimulation on cultured myoblasts and understanding the role of the ECM in migration, proliferation and differentiation of the cells [453,[456], [457], [458], [459], [460]]. Another approach in developing a higher differentiation and more functional skeletal muscle tissue is the application of electrical stimulation [[461], [462], [463]]. This stimulus is important in inducing the appropriate functions of native skeletal muscle. Similar to nerve stimulation that occurs during myogenesis and regeneration of injured tissue, electrical stimulation has shown to induce differentiation of neural activity [452], where the primary myoblasts can influence the expression of a myosin heavy chain [464] and myogenic transcription factors [465], while also promoting maturation of 2D primary muscle cultures in vitro [462,466], and even increase cell proliferation [467].
Piezoelectric materials have demonstrated considerable potential for skeletal TE (Table 8). Martins and co-workers studied the influence of polarization and morphology of PVDF films and fibers, on the adhesion, morphology and proliferation of a mouse myoblast cell line, with the aim of determining the optimal environmental conditions required for muscle regeneration. The authors reported that negatively poled charges improved cell adhesion and proliferation, which was synergistic with oriented fibers that supported directional growth of these cells [468].Table 8Applications of piezoelectric materials in skeletal and cardiac tissue engineering.Table 8MaterialValidation in vitroValidation in vivoReferencePVDF films/fibersNegatively poled improved adhesion, proliferation; oriented fibers directional growth in mouse myoblasts–[468]ZnO nanorods on PDMS with PNIPAAmInduced myogenic differentiation in hUCB-MSCsImproved retention, survival, regeneration in mouse injury[469]ZnO nanogeneratorsProliferation, differentiation in skeletal/smooth muscle; Ca increase in smooth only–[470]Bio-based diacid/diol elastomerInduced myogenic differentiation, improved morphology/performance–[471,472]PVDF-TrFE scaffoldsAttachment, survival, alignment, sarcomeres, contractile, minimal apoptosis in stem-derived cardiomyocytes–[482]PCL nanofilm with PVDF-TrFE microfibersMimicking ECM for cardiovascular cells–[483]PVDF-TrFE electrospunDeveloped functional cardiac tissues from cardiomyocytes; sensors for contraction–[484]CoFe2O4-PVDFImproved proliferation, spread in oxidative stress cardiomyocytes–[485]PVDF with bilayer grapheneDifferentiated H9c2 to cardiomyocyte, activated signaling–[486]ZnO nanorods multilayerEx vivo cardiomyogenic differentiation, alignment, expression in hMSCs–[487]PCL/PVDF/PCL-Fe3O4Maturation, structure, function, gene in cardiomyocytesRepaired ventricle, enhanced EF, reduced inflammation/fibrosis in rat MI[488]
In another study, Yoon and collaborators developed a two-step strategy for skeletal muscle regeneration in mice. First, they employed a piezoelectric substrate, made of aligned ZnO nanorods spin-coated with polydimethylsiloxane followed by surface grafting with poly(N-isopropylacrylamide), to serve as a mechanoelectrical inducer of myogenic differentiation of human umbilical cord blood mesenchymal stem cells (Fig. 6b). In the subsequent step, cell sheet fragments derived from these myogenic cells were harvested from the piezoelectric substrate and injected into injury sites of mouse skeletal muscle. This approach resulted in improved cell retention, enhanced cell survival, and overall muscle regeneration [469].Fig. 6Applications of piezoelectric materials in skeletal and cardiac muscle tissue engineering. a) A piezoelectric elastomer was fabricated through random copolymerization to achieve low elastic modulus and stretchability similar to that of skeletal muscle. In vivo implantation and stimulation promote myogenic differentiation, muscle repair and functional recovery. Reproduced with permission from Ref. [472]. b) a thermosensitive, stretchable and piezoelectric substrate serves as a platform for myogenic differentiation of hUCBMSC cells. Reproduced with permission from Ref. [469]. c) A biocompatible piezoelectric elastomer composed of bio-based diacids and diols matches the elastic modulus of muscle, provides in situ electrical stimulation and promotes myoblast proliferation and differentiation and improves muscle repair. Reproduced with permission from Ref. [471]. d) A contactless dynamic culture platform with magnetic and piezoelectric stimulation drives cardiomyocyte maturation for PCL-PVDF cardiac constructs, that are e) implanted in infarcted rat hearts and generate voltage output response. Reproduced with permission from Ref. [488]. f) A self-powered CoFe2O4-PVDF composite nanofiber sheet serves as a cardiac patch. An external magnetic field enhances the piezoelectricity of the nanofiber sheet promoting steroid hormone biosynthesis, opening of potassium and calcium ion channels, contraction of injured cells and repair of heart failure. Reproduced with permission from Ref. [485]. g) A multilayer stretchable PDMS-Zinc oxide composite serves as a platform for application of pulsatile mechanoelectric cues that stimulate cardiomyogenic differentiation of hMSCs ex vivo. Reproduced with permission from Ref. [487].Fig. 6
In a later study, ZnO nanosheets were fabricated on glass coverslips via the deposition of an aluminum nitride (AlN) followed by immersion in a solution of hexamethylenetetramine and Zn(NO3)2 within a sealed jar in an oven at 80 °C for 9 h. The resulting ZnO nanogenerators produced a piezoelectric effect that was able to induce proliferation and differentiation in both skeletal and smooth muscle cell while maintaining high cell viability. Interestingly, the electromechanical cues produced by these ZnO nanogenerators were able to stimulate smooth muscle cells, by increasing calcium levels in the cytoplasm, but did not exert the same effect in skeletal muscle cells. This difference was attributed to the random orientation of the ZnO nanogenerators, underscoring the importance of alignment for effective stimulation of skeletal muscle cells [470].
More recently, a promising strategy for treating skeletal muscle loss focused on the development of a biocompatible and biodegradable piezoelectric elastomer by copolymerization of bio-based diacid and diol. By tuning the content of the elastomer, these authors were able to obtain a scaffold with elastic modulus matching that of native muscle tissue, allowing it to accommodate tissue movement. Furthermore, the scaffold generated a piezoelectric charge that induced myogenic differentiation, and improved the morphology and performance of the skeletal muscle (Fig. 6a–c) [471,472].
In the human heart cardiomyocytes contract synchronously, due to an electrical conduction system, that is vital in maintaining correct heart function. Pacemaker cells in the sinoatrial node are the generators of electrical impulses, which then propagate through the atrioventricular node and Purkinje fiber network to elicit contraction of ventricular cardiomyocytes. Intercellular communication structures allow rapid electric signal propagation and maintain cellular cohesion, transmitting mechanical force to neighbor cardiomyocytes, resulting in rhythmical and synchronized contraction and relaxation of the cardiac muscle [473].
Cardiac muscle tissue engineering envisages the reconstruction of a microenvironment that mimics the complex biophysical and biochemical cues that promote the organization of myocardial cells in a functional manner, to repair or replace damaged heart tissue [473,474]. Studies of cardiac tissue engraftment in animals models have shown improved cardiac function and attenuated remodeling after myocardial remodeling [475]. Additionally, several works have demonstrated that the application of external electrical stimulation can replicate the natural electrical signals of the myocardium and promote its propagation in myocardial cells, while also supporting tissue organization and synchronized contraction [474]. Moreover, the incorporation of conductive scaffolds with electrical fields contributed to recreate a native microenvironment, by promoting cardiac tissue maturation, cell proliferation, cardiomyogenic differentiation and expression of cardiac markers [[476], [477], [478], [479]]. Notably, Shin et al. [480] achieved 3 times higher spontaneous synchronous beating rates and 85 % lower excitation threshold in neonatal rat cardiomyocytes cultured on carbon nanotube-incorporated photo-cross-linkable gelatin methacrylate GelMA hydrogels. In another elegant study, Feiner et al. [481] developed a cardiac patch that included a porous electronic mesh with multiple electrodes. This strategy allowed for multiple functionalities, such as providing electrical stimulation, releasing of biochemical factors to the engineered tissue and recording tissue function.
In consequence of the observations of improved cardiac cell response to scaffolds that favor electric signal propagation, new strategies were developed. On the basis of the electroactive nature of the myocardium and that scaffolds that combine mechanical and electrical properties may foster the development of a functional cardiac tissue, the application of piezoelectric materials in cardiac muscle tissue engineering has recently gained traction.
The effect of piezoelectric PVDF-TrFE scaffolds on stem cell-derived cardiomyocytes was studied by Hitscherich et al. [482]. Cardiovascular cells were shown to attach to and survive with minimal apoptosis when cultured in the scaffolds. Additionally, high alignment was observed, as well as defined sarcomeres, and the scaffolds supported spontaneous contractile behavior. Similar results were observed by Gouveia et al. [483] that developed a scaffold based on a PCL nanofilm with magnetic properties coated with a layer of piezoelectric PVDF-TrFE microfibers with the purpose of mimicking cardiac extracellular matrix. PVDF-TrFE was also employed to develop dual function scaffolds [484]. By electrospinning PVDF-TrFE copolymers authors obtained increased β phase resulting in increased piezoelectricity of scaffolds. Authors showed that the scaffold can be used as a platform to develop functional cardiac tissues from cardiomyocytes. Additionally, scaffolds can function as sensors translating the contractile activity of the engineered tissue to readable electric signals. An alternative strategy to use PVDF in cardiac patch development involved its doping with magnetostrictive CoFe2O4 [485]. This resulted in increased piezoelectric ratios due to the increase in β phase and crystallinity. Via electrospinning authors developed CoFe2O4-PVDF composite nanofiber sheets that were tested for the culture of oxidative stress injured cardiomyocytes. The proliferation and spread area of injured cells was shown to improve when co-cultured with CoFe2O4-PVDF and receiving magnetoelectric cues (Fig. 6f). In another study [486] PVDF nanofibers were modified with bilayer graphene nanolayers. Researchers, aimed at combining the high electroconductivity of graphene with the piezoelectricity of PVDF to develop a stimuli-responsive biomaterials. When cultured in the scaffolds, H9c2 cardiomyoblasts successfully differentiated towards cardiomyocyte phenotype. Results shown that pulsatile electrical stimulation induced activation of β1-integrin-mediated signaling pathways and opening voltage gated calcium channel. Yoon et al. [487] opted to use ZnO nanorods to develop a stretchable piezoelectric substrate with the aim of providing pulsatile mechano-electric cues. The substrate consisted in a multistacked layer-by-layer composite composed of aligned ZnO nanorods on a PDMS substrate. Cell culture on the substrate resulted in ex vivo cardiomyogenic differentiation of hMSCs, accompanied of cell alignment and increased expression of cardiac transcription factors, and structural, ion channel and gap junction proteins (Fig. 6g).
Overall, the diversity of piezoelectric scaffolds proposed shows high potential to engineer functional cardiac muscle tissue. Nevertheless, the biocompatibility and functionality of piezoelectric based cardiac scaffolds still requires in vivo testing. In this regard, Han et al. [488] recently developed a triple material scaffold. Authors employed a melt-based EHD printing process to deposit serpentine 10.13039/100018919PCL fibers, serving as mechanical support. Then a solution-based EHD printing process was used to deposit serpentine PVDF fibers, serving as piezoelectric microenvironment generator. Finally, an extrusion printing process was used to deposit magnetic PCL/Fe3O4 fibers at one of the extremities of the scaffold to induce deformation of the engineered tissue. A moving magnet was then employed to interact with PCL/Fe3O4 fibers to generate a driving force that was shown to promote in vitro maturation of the cardiomyocyte phenotype, accompanied with improved cellular structure, function and gene expression (Fig. 6d). Authors, further tested the cardiac constructs in vivo on myocardial infarction rat models by transplantation on the infarcted area (Fig. 6e). The engineered cardiac constructs showed encouraging potential in repairing the damaged ventricle. Improved myocardial response was demonstrated by enhanced ejection fraction and suppressed inflammation and fibrosis was observed, underscoring promising therapeutic potential of piezoelectric engineered cardiac tissue for myocardial infarction therapy.
Although additional research is needed to optimize the production of piezoelectric structures for skeletal and cardiac muscle TE, the studies here presented collectively underscore the critical importance physical and electrical stimuli for proper in vivo muscle development. Ultimately, fine-tuning piezoelectric material properties will be key to eliciting improved cell responses closely mimicking the native niche, and achieving successful muscle regeneration.
The skin is the largest organ of the body and a physical barrier protecting the internal body systems from the outside environment. When injured, highly orchestrated intra- and intercellular pathways are activated to restore tissue integrity [489]. However, aberrations normally associated to disruption in the remodeling phase of physiologic wound healing can lead to pathologic scarring and leading to tissue dysfunction. This abnormal healing has been largely linked to the cellular elements within granulation tissue – fibroblasts and myofibroblasts – that generate contractile forces within the wound milieu, leading to direct inward movement of wound surroundings through complex intracellular and extracellular interactions [490]. Such wound dynamics causes simultaneous alteration of physical properties of the tissues alongside ongoing biological transformations of the micro and macroenvironments, a process known as biophysical coupling.
Injury that disrupts an epithelial layer instantaneously generates endogenous electric fields. These fields, which have been proposed as a directional cue that guides cell migration in wound healing [[491], [492], [493], [494]]. Several studies have explored the mechanism by which cells may perceive this electric potential [336,[495], [496], [497]]. For example, McLaughlin and Poo [497] suggested that the electric currents exerted forces on freely diffusing signaling molecules on the extracellular surface leading to an asymmetric redistribution of these molecules and hence, directed migration. In accordance with what happens in bone [344,345], Robinson [495] proposed that electric fields would induce asymmetric activation of a plasma membrane voltage sensor, such as a voltage-gated calcium channel, or other transmembrane voltage sensor [336,496]. Asymmetric activation of this voltage sensor triggers a polarizing signal that directly affects cell morphology other cellular responses [498].
In addition to endogenous electric field, the exertion of mechanical stress to the healing wound produces microdeformations in collagen fibers, which increase the local negative electrical charges in situ through piezoelectric effect [499]. The concentration of localized negative charges acts as stimulatory cue to enhance proliferation of soft tissue cells in the vicinity of the mechanically stressed regions, through the piezoelectric phenomenon. Thus, both the quality and quantity of piezoelectric current may change concurrently reflecting dynamic changes in tissue properties as the healing process progresses. It appears that mechanical stress regulates not only wound contraction, but also influences ECM remodeling by modulating the biological-physical-mechanical axis. In this interplay, mechanical stimulus affects the production of piezoelectricity, which in turn modulates the expression and signaling of transforming growth factor-β (TGF-β). Subsequently, this growth factor affects the production of piezoelectricity through alteration of ECM properties, such as the mechanical characteristics of collagen fibers [500]. Therefore, the modulation of biological-physical-mechanical axis can have impact on wound contraction and resultant scarring, while the use of exogenous electrical stimulation may mimic the endogenous current and improve wound healing outcomes.
Farahani and Kloth [339], proposed exploiting of the converse piezoelectric effect as a therapeutic modality to diminish scar-related contraction by harnessing the skin's intrinsic piezoelectric properties. De Rossi et al. [335] compared two synthetic polymer analogs with the aim at providing further insight into the nature of the piezoelectric activity of true epidermis. The first model was made of a film of uniaxially oriented PHB, whilst the second analog was made of a PVDF film. Comparative evaluation of these two polymers have shown a striking similarity between the piezoelectric properties between characteristics of PHB and those of epidermis. In a subsequent study, PHBV was combined with polydioxane (PDX) to take advantage of the piezoelectric properties of PHBV (comparable to that of collagen) and PDX's mechanical properties (similar to elastin and collagen). These authors reported a combination of 80/20 PHBV/PDX scaffold was found to support cell migration, proliferation and adhesion while also promoting angiogenesis, keratinocyte proliferation and reduced inflammation-factors that accelerated skin regeneration in rats [501].
Other approaches have utilized piezoelectric electrospun PVDF-TrFE fibers (Table 9). Weber et al. [502] reported on in vitro behavior of human skin fibroblasts cultured on these fibers and compared to cells grown on tissue culture polystyrene. Results demonstrated that cell viability, proliferation and morphology of cells cultured on both templates were comparable over a 7-day period although the direct influence of piezoelectric effects on the cultured cells was further investigated. In 2018, Wang et al. [503] showed that tweaking electrospinning parameters can optimize the piezoelectric properties of PVDF-TrFE nanofibers. These scaffolds induced perfect alignment of fibroblasts along the electrospinning direction and an increase in cell proliferation rate. Electrospun PDFV-TrFE scaffold properties were further enhanced by addition of tetrabutylammonium chloride (PVDF-TrFE/TBAC), that resulted in 5.3 fold higher voltage output [504]. Under mechanical stimulation, these membranes promoted cell growth and accelerated wound healing by 37 %, in addition, to exhibiting a strong antibacterial effect.Table 9Applications of piezoelectric materials in skin tissue engineering.Table 9MaterialValidation in vitroValidation in vivoReferencePHBV/PDX 80/20Supported migration, proliferation, adhesionPromoted angiogenesis, keratinocyte proliferation, reduced inflammation, accelerated regeneration in rats[501]PVDF-TrFE fibersViability, proliferation, morphology comparable in human fibroblasts–[502]PVDF-TrFE nanofibersInduced alignment, increased proliferation in fibroblasts–[503]PVDF-TrFE/TBACPromoted cell growthAccelerated healing by 37 %, strong antibacterial[504]PU/PVDF fibersEnhanced NIH3T3 migration, growth, proliferationAccelerated healing, doubled speed[505]PVA/PVDF hydrogel–Induced re-epithelization, vascularization, ECM, anti-inflammatory, accelerated healing in rat diabetic[506]Aligned PVDF with polydopamine-polyacrylamideIncreased vascularization, ECMHair follicle formation, accelerated healing[507]Ag-PVDF with PCL/gelatin JanusPromoted fibroblast proliferation, migrationAccelerated healing, antibacterial, drainage in mice[508]ZnO nanorods on PDMS–Increased collagen, proliferation, re-epithelization[509]ZnO/PVDF/SA hydrogel–Accelerated healing in rat diabetic[510]ZnO in CA matsImproved adhesion, proliferation, migration–[511]Polydopamine on chitosan film–Higher regeneration, proliferation, angiogenesis, collagen[512]Silk fibroin with LiNbO3/CNT–Improved healing in mouse[513]BaTiO3 with PEG/PVDF-TrFE/ZIF-8/goldHigh antibacterial efficiency; promoted fibroblast proliferation/migration, collagenAccelerated wound healing in vivo[[514], [515], [516], [517], [518]]CM-chitosan/tannic/carbomer/FeWO4High antibacterial; promoted proliferation migration collagenAccelerated healing[519]PLLA nanofibersHigh antibacterial; promoted proliferation migration collagenAccelerated healing[520]
PVDF has also been used in conjunction with other materials. For example, electrospun polyurethane/PVDF (PU/PVDF) fibers were developed with the aim of studying the piezoelectric effect on NIH3T3 cells activity in vitro and in vivo [505]. The results showed enhanced cell migration, growth and proliferation when compared to both paraelectric fibers and PU fibers under dynamic stimulation, approximately doubling the wound healing speed, as compared to on controls. Similar effect was observed in vivo, were piezoelectricity induced by random animal movements resulted in accelerated wound healing.
Wang and collaborators combined poly vinyl alcohol (PVA) and PVDF (PVA/PVDF) to produce a piezoelectric hydrogel via a process of annealing-swelling that enhanced the hydrogen bonding between PVA and PVDF while obtaining an electroactive β-phase and crystalline structure [506]. This enabled the construct to uniformly convert the mechanical load from movement into a piezoelectric stimulus. In rat diabetic wound model, the hydrogels induced significant re-epithelization, vascularization and ECM secretion. This affected the macrophage phenotype resulting in an anti-inflammatory environment that accelerated healing (Fig. 7b).Fig. 7Applications of piezoelectric materials in skin tissue engineering. a) A PVDF/sodium alginate scaffold was modified with zinc oxide nanoparticles. The scaffold has dual piezoelectric release in vertical and horizontal directions that stimulates cell proliferation and migration, collagen deposition and neovascularization resulting in accelerated wound healing. Reproduced with permission from Ref. [510]. b) A self-powered piezoelectric PVA/PVDF composite hydrogel was implanted in a diabetic wound rat model promoting regulated inflammation, growth factor secretion, angiogenesis and cell migration and proliferation, thereby promoting wound healing. Reproduced with permission from Ref. [506]. c) A piezoelectric dermal patch based on bidirectionally grown zinc oxide nanorods aligned on PDMS. A PEDOT:PSS layer was used as antiadhesive layer between the PDMS and the glass base. A one-directional rubbing process was used to align nanorods. The monolayer of zinc oxide nanorods was spin-coated with PDMS. The repetition of the process resulted in multilayer patches. Reproduced with permission from Ref. [509]. d) A polydopamine coated chitosan film generates piezoelectricity under mechanical pressure an presents photothermal effect when irradiated with near-infra red light. In in vivo full thickness wounds, the film enhances wound healing. Reproduced with permission from Ref. [512]. e) A nanofibrous scaffold, composed of silk fibroin, LiNbO3 and multiwalled carbon nanotubes was fabricated by electrospinning for wound healing application. Reproduced with permission from Ref. [513]. f) A self-powered piezoelectric and biodegradable wound healing dressing was developed based on PLLA, PEG and tetragonal barium titanate. The piezoelectricity of the dressing eliminates bacteria and improve wound healing rate. Reproduced with permission from Ref. [515]. g) An ultrasound-controlled nanocomposite was developed based on self-assembly of zeolitic imidazole framework-8 on the surface of barium titanate loaded with ciprofloxacin. Ultrasound activation of the nanocomposite enhances reactive oxygen species generation creating an antibacterial environment and improves wound healing rate. Reproduced with permission from Ref. [517]. h) A biodegradable self-charged piezoelectric PLLA nanofiber matrix. Activation by external ultrasounds produces controllable surface charges, prevents bacterial infection and promotes wound healing. Reproduced with permission from Ref. [520].Fig. 7
Du et al. [507] proposed combining aligned PVDF nano fibers with a polydopamine-polyacrylamide hydrogel. The fiber alignment increased the piezoelectric potential of the patches while the hydrogel composition improved self-adhesion properties. The patch increased the expression of markers of vascularization and ECM deposition was observed both in vivo and in vitro, and even partially stimulated hair follicle formation. Aligned Ag nanoparticles-doped PVDF nanofibers were employed in wound dressings, acting as a hydrophobic layer, while the hydrophilic layer was composed of polycaprolactone/gelatin nanofibers [508]. These Janus dressings proved successful in converting mechanical stimuli from body movements into electrical energy. This piezoelectric effect promoted fibroblast proliferation and migration, and accelerated the rate of wound healing in mice, while also reducing bacterial infection and facilitating the drainage of wound exudate.
Zinc oxide (ZnO) nanorods, have been used in wound patches as the base to capture mechanical movement and transform it into an electrical field. Bhang et al. [509] developed a patch of ZnO nanorods on a PDMS layer. Results indicated that by aligning the ZnO nanorods and increasing its density, the piezoelectric properties of the patches were improved (Fig. 7c). The voltage generated enhanced the wound healing process by increasing collagen deposition, cell proliferation and re-epithelization. Zengjie Fan's group [510] employed 3D printing technology to fabricate a piezoelectric hydrogel composed of ZnO nanoparticles modified PVDF and sodium alginate (SA). This approach produced a dual piezoelectric release model where SA absorbed wound exudate, creating vertical swelling and horizontal friction that provided permanent electrical stimulation and acceleration of the wound healing process (Fig. 7a) [510]. Ghosh et al. [511] incorporated up to 1.5 % ZnO into cellulose acetate (CA) fiber mats in order to enhance its piezoelectric performance while reinforcing tensile strength of the electrospun mats and maintaining its degradability. In vitro assessment showed improved cell adhesion and proliferation, and faster migration in wound closure assays.
Chen et al. [512] developed a dual functional film with both piezoelectric and photothermal properties, by coating polydopamine onto a chitosan film. The film generated electric voltages in response to mechanical stimuli with higher piezoelectric voltage increased with increased dopamine concentration and also with near-infra red irradiation (NIR). In an in vivo wound model, these authors demonstrated that the films combining chitosan and polydopamine and receiving NIR promoted higher wound regeneration potential by inducing cell proliferation, angiogenesis and collagen deposition (Fig. 7d). Additionally, Yue et al. [513] demonstrated that silk fibroin nanofibers alone are non-piezoelectric. However, when combined with LiNbO3 nanoparticles and carbon nanotubes (LN/CNT/SF-NFS), the resulting scaffolds responded to pressure stimulation generating piezoelectric voltage that was stable for 200 cycles of pressure-release. Furthermore, when applied on mouse wounds the healing process was significantly improved (Fig. 7e).
In recent years US have emerged as an alternative form of generating piezoelectric current, particularly in wound repair strategies, particularly those aimed at combating the dual treatment of bacterial wound infection and skin tissue regeneration. In most cases, researchers have developed piezoelectric nanocomposites based on the use of BaTiO3 combined with other materials such as poly(ethylene glycol) (Fig. 7f) [514,515], PVDF-TrFE [516], zeolitic imidazolate framework-8 (Fig. 7g) [517] or gold particles [518]. Xu and coworkers developed an expandable and degradable hydrogels that were assembled along with carboxymethyl chitosan, tannic acid, carbomer and FeWO4 nanorods [519], while Das and collaborators opted for PLLA nanofiber matrices (Fig. 7h) [520]. In these approaches the piezoelectric scaffolds are activated by external US which trigger the release of reactive oxygen species, that showed high antibacterial efficiency. At the same time the sonodynamic process promoted fibroblast proliferation and migration, collagen deposition, and accelerated the process of wound healing in in vivo models.
The piezoelectric phenomenon, as discussed in previous sections, originates from the electromechanical coupling in non-centrosymmetric materials, enabling the conversion of mechanical stress into electrical signals and vice versa. While earlier parts of this review focused on the theoretical foundations and synthetic materials like PVDF, recent advancements have illuminated cross-scale mechanisms that bridge molecular-level processes in polymers to macroscale responses in human tissues. These innovations emphasize how nanoscale structural asymmetries in piezoelectric polymers can generate electrical cues that mimic and influence biological piezoelectricity, fostering applications in tissue engineering. This section explores these mechanisms, drawing from peer-reviewed studies that demonstrate scalable transitions from polymer chain orientations to tissue-level regeneration, highlighting self-powered systems that enhance cellular behaviors without external energy sources.
At the molecular scale, piezoelectricity in polymers arises from dipole moments induced by asymmetric chain conformations, which can be engineered to amplify electromechanical responses. For instance, in PVDF and its derivatives, the β-phase conformation features a zigzag arrangement of carbon-fluorine and carbon-hydrogen bonds, creating net polarization under mechanical deformation. Recent work has advanced this through interface engineering techniques. One study introduced a liquid-liquid interface polar engineering method to induce polar asymmetry in ultra-soft organic films composed of polyethylene glycol and polystyrene-block-polyisoprene-block-polystyrene PEG/SIS [521]. This approach overcomes steric hindrance by aligning polar groups at the interface, resulting in a piezoelectric coefficient comparable to PVDF, while maintaining ultra-softness akin to human skin or cartilage. The mechanism involves solvent-induced orientation during film formation, where water and toluene create a polar gradient that stabilizes asymmetric layers. This cross-scale innovation extends from nanoscale dipole alignment, confirmed via density functional theory (DFT) calculations showing enhanced interactions between polyethylene glycol hydroxyl groups and SIS, to microscale film uniformity, enabling large-scale production with consistent piezoelectric output. Such materials demonstrate how molecular tailoring can propagate to tissue-compatible softness, supporting biomechanical sensing in vivo without rigidity mismatches.
Building on these polymer innovations, cross-scale mechanisms reveal how nanoscale electrical signals influence cellular and tissue-level processes, mirroring innate piezoelectricity in human tissues. Biological piezoelectricity, observed in collagen-rich structures like bone and tendon, stems from non-centrosymmetric helical conformations that generate charges under stress, facilitating remodeling. At the molecular level, collagen fibrils in bone exhibit piezoelectricity, where mechanical compression displaces charges, altering local electric fields [329]. These fields then scale to cellular interactions, activating ion channels and signaling pathways such as Wnt/β-catenin, which promote osteoblast proliferation and matrix deposition. To validate this in practice, an original experimental study developed piezotopographically engineered scaffolds by dispersing HAp nanoparticles with PVDF-TrFE with aligned topography to enhance bone regeneration [522]. In vitro experiments with hMSCs showed that the scaffolds, led to an increase in alkaline phosphatase activity and calcium deposition compared to non-piezoelectric controls. In vivo implantation in rat calvarial defects demonstrated accelerated bone formation, after 6 weeks. This mechanism bridges scales by converting polymer-generated microvolt potentials into tissue-level responses, such as improved ECM synthesis.
In neural tissues, cross-scale piezoelectric mechanisms have been innovated to guide regeneration through oriented electrical gradients. Piezoelectric polymers like PVDF-TrFE create self-powered scaffolds that respond to mechanical strains, generating potentials that direct axon growth. This nanoscale effect scales to microscale hydrogel networks, where embedded nanorods form porous structures that mimic neural ECM, enhancing mechanotransduction. Complementing this, Tai et al. [523] fabricated electrospun conductive piezoelectric scaffolds using PVDF-TrFE for peripheral nerve repair. In vitro tests with Schwann cells under dynamic compression showed enhanced cell proliferation and upregulated expression of neurotrophic factors like Ngf. In vivo in a rat sciatic nerve defect model, the scaffolds promoted axonal regeneration, with increased myelinated fibers. These findings illustrate how polymer-derived nanoscale charges influence macroscale neural connectivity. Another study by Shi et al. [524] explored a biomimetic piezoelectric hydrogel for spinal cord injury, incorporating KNN nanoparticles with decellularized spinal cord matrix gel. In vitro with neural stem cells, ultrasound-triggered piezoelectricity increased ATP production via enhanced mitochondrial function. Using a rat spinal cord hemisection model, authors demonstrated improved differentiation of neural stem cells into functional neurons.
Similar cross-scale innovations apply to skin and muscle regeneration, where piezoelectric polymers adapt to dynamic mechanical environments. In skin tissue engineering, polyurethane/PVDF composites generate microelectric fields under deformation, stimulating fibroblast migration and collagen deposition, serving a crucial role in modulating immune responses for wound healing [525]. In a recent study Liu et al. [526] developed a piezoelectric-immunomodulatory electrospun membrane using PLLA doped with gallium-mesoporous bioactive glass. In vitro assays with macrophages under lipopolysaccharide stimulation showed that the membrane, activated by ultrasound, reduced pro-inflammatory cytokine IL-1β and increased anti-inflammatory IL-10, while promoting angiogenesis. In diabetic mouse wound models, treatment accelerated wound closure and enhanced collagen deposition and vascularization. For skeletal muscle, oriented PVDF fibers align myoblasts directionally, with nanoscale fiber topography amplifying piezoelectric output. Xu el al [527]. fabricated electrospun PVDF-TrFE/ceria nanoparticle scaffolds for volumetric muscle loss. Testing with C2C12 myoblasts, the scaffolds under cyclic strain induced higher myosin heavy chain expression indicating myotube formation. In mouse tibialis anterior defects, the composite scaffolds restored grip strength faster and increased muscle fiber diameter with reduced fibrosis.
In neurodegenerative applications, recent trends emphasize composite piezoelectric materials that span scales for diagnosis and therapy. Organic polymers like PLLA, combined with inorganic fillers, convert mechanical energy into electrical signals at the nanoscale, enabling sensitive biosensors for amyloid detection in Alzheimer's disease [528]. The cross-scale mechanism involves nanoparticle dispersion creating localized fields that enhance protein binding specificity, with surface plasmon resonance data validating detection limits down to picomolar levels. At the tissue scale, these materials form implantable devices that stimulate neural repair, as seen in ALS models where piezoelectric stimulation improved motor function scores.
These recent innovations highlight the promise of cross-scale piezoelectric mechanisms in bridging synthetic polymers to human tissues. By engineering molecular asymmetries, researchers can generate scalable electrical signals that direct cellular behaviors, offering dynamic, self-sustaining platforms for regenerative medicine. Future work may focus on multimodal composites integrating piezoelectricity with other stimuli, potentially revolutionizing treatments for complex tissue defects.
The integration of piezoelectric properties into composite materials has opened new avenues for bionic designs that emulate the electromechanical coupling observed in natural tissues. These bioinspired composites leverage the inherent piezoelectric responses of biological structures, such as the aligned collagen fibrils in bone or the dynamic signaling in neural interfaces, to create scaffolds that respond to physiological mechanical cues. These composites typically combine piezoelectric polymers like PVDF or PLLA with ceramic fillers such as BaTiO3 to enhance charge generation and mechanical stability. By combining polymers with ceramic or molecular fillers, researchers have developed materials that not only provide structural support but also deliver localized electrical stimulation to guide cellular processes.
One recent strategy is to impose structural gradients or multi-scale architectures that mirror the organization of natural load-bearing tissues. For example, a recent DLP 3D printing study fabricated BaTiO3/β-TCP scaffolds with a triply periodic minimal surface (TPMS) architecture and spatial variation of ceramic loading. Results showed that increasing BaTiO3 content increased the local d33 response, while the scaffold structure remained permeable and mechanically viable [529]. In this way, the spatial modulation of filler content yields a composite whose piezoelectric performance is co-tuned with porosity and mechanical stiffness. In a different direction, an adaptive mineralization concept was proposed in which a composite scaffold whose internal piezopotential triggers localized mineral deposition from the surrounding ionic milieu in proportion to applied stress [530]. The authors showed in a model system that piezoelectric charges preferentially drive mineral nucleation in higher stress zones, resulting in a self-reinforcing adaptation of mechanical stiffness over time. Although still conceptual, this illustrates how bionic feedback loops might be built into composite designs.
To reconcile the mechanical mismatch between stiff piezoceramics and soft tissues, many groups are developing composite scaffolds that embed piezoelectric ceramic fillers into flexible, often biodegradable polymer matrices. But a key innovation is the more intelligent coupling between phases to enhance effective electromechanical coupling and interfacial stress transfer. Barkow et al. [531] used direct ink writing of PCL/BaTiO3, and BaTiO3 plus bioactive glass composites to fabricate scaffolds. In this case, using 40 vol % BaTiO3 yielded a d33 comparable to bone, while adding bioactive glass improved osteoblastic response, even though it slightly reduced piezoresponse. This demonstrates a trade-off sacrificing some piezopotential for better bioactivity, but within a bionic composite framework. Another recent innovation is the design of a piezoelectric and conductive nerve conduit using electrospun nanofibers combining piezoelectric polymers and conductive fillers to self-generate electrical stimulation under deformation. The authors followed a synergistic approach combining piezoelectric and conductive phases in a fiber geometry reminiscent of nerve fascicles, demonstrating that the composite conduit could support neurite extension and guide regeneration without an external electrode [532].
A key requirement for clinical translation is that piezoelectric scaffolds degrade or resorb in a biologically benign way after fulfilling their regenerative role. Recent work has explored composites in which at least one piezoelectric phase can be enzymatically digested or resorbed, while the remaining scaffold maintains structural integrity during tissue regrowth. In a 2025 study, Jarkov et al. [533] fabricated composite films and scaffolds of KNN integrated into cellulose, with the cellulose component subject to enzymatic degradation. With this design authors leverage a bioresorbable ‘sacrificial’ matrix integrated with a piezoceramic for transient function. Resutls demonstrated that the composite still exhibits measurable piezoresponse, promotes neural stem cell differentiation and is gradually degraded under cellulase exposure.
A further bionic twist is using non-contact, remotely delivered mechanical waves, such as US, to activate piezoelectric composites in vivo, mimicking how certain biological systems respond to mechanical perturbations, for example bone vibration. Several recent studies used ultrasound to stimulate polymer/ceramic piezoelectric composites in situ. Recently, Chen et al. [370] implanted hierarchically architected PVDF/ZIF-8 foams in rat femur defects and activated by ultrasound, resulting in enhanced angiogenesis and osteogenesis. The authors reported upregulation of ion transport and improved bone metrics in vivo compared with non-stimulated controls. Similarly, a PVDF-TrFE/BaTiO3 nanoparticle composite was shown to mediate Ca^2+^ transients and promote differentiation of SH-SY5Y cells under ultrasound stimulation [534]. These works exemplify how bionic design now includes the stimulus mode as part of the composite system.
A compelling trend is modular composites that embed distinct functional a piezoelectric module for electro-stimulation, a bioactive module for osteoinduction, and an immunomodulatory module. Some recent works show that by spatially partitioning these functionalities, the composite mimics the multifunctionality of natural bone. In this regard, Chen et al. [535] evaluated the immunomodulatory response of macrophages to a piezoelectric/ceramic composite scaffold. The piezoelectric activation was found to shift macrophage polarization toward an M2 phenotype, while also stimulating osteogenic gene expression in co-cultured stem cells.
Although recent strategies in the bionic design of piezoelectric composites have demonstrated considerable promise, a number of limitations remain that must be resolved before clinical translation can be achieved. One of the central issues concerns the interfacial mechanics between soft polymeric matrices and rigid ceramic fillers. The generation of a stable piezoelectric output depends on efficient stress transfer across these interfaces; however, delamination and microcracking under cyclic loading continue to compromise long-term functionality. A second limitation arises from the inherent trade-offs between electromechanical response, mechanical stability, and biological activity. Composites that exhibit high piezoresponse frequently do so at the expense of structural robustness or bioactivity, thereby necessitating approaches that integrate multi-objective optimization rather than focusing on a single property. A further challenge relates to the behavior of these systems under physiological conditions. Many of the recently developed composites are designed to be partially degradable or resorbable, yet the timing of this degradation is often not synchronized with the natural processes of tissue healing, resulting in premature loss of piezoelectric functionality. Finally, questions of scalability remain significant. Several of the most advanced architectures rely on additive manufacturing techniques or precise control of filler distribution, methods that are effective at a laboratory scale but are not yet proven at the dimensions required for clinical use. Taken together, these factors highlight that while bionic design has already opened new pathways for mimicking the hierarchical and multifunctional nature of biological tissues, further refinement is needed to achieve robust, reproducible, and clinically relevant systems. Nevertheless, the incorporation of bioinspired concepts into composite design provides a compelling strategy for narrowing the gap between synthetic materials and the adaptive complexity of living tissues, and it is expected to play an increasingly important role in the development of the next generation of piezoelectric scaffolds for tissue engineering.
The exploitation of piezoelectricity in tissue engineering has gained significant traction in recent years with experimental findings underscoring the tissue-specific impacts of piezoelectric polymers (Fig. 8). In bone tissue engineering, piezoelectric scaffolds, such as those incorporating PVDF or BaTiO3 composites, have demonstrated enhanced osteoblast adhesion, proliferation, and mineralization. For instance, under mechanical loading, these materials generate surface charges that promote osteogenesis by activating calcium signaling pathways and upregulating osteogenic genes like RUNX2 and osteocalcin, leading to accelerated bone defect repair in animal models. Similarly, in cartilage regeneration, biodegradable piezoelectric polymers like PHBV and injectable hydrogels with piezoelectric nanoparticles have been shown to facilitate chondrocyte migration and extracellular matrix deposition. Ultrasound-activated stimulation from these scaffolds enhances TGF-β expression and calcium influx, resulting in improved neocartilage formation and reduced osteoarthritis progression in rabbit models. For nerve tissue engineering, piezoelectric conduits made from PVDF-TrFE or chitosan composites promote neurite outgrowth and Schwann cell alignment. The generated electrical fields under dynamic loading stimulate axonal regeneration and myelination, as evidenced by faster functional recovery in rat sciatic nerve defect models, potentially through modulation of voltage-gated ion channels and neurotrophic factor release. In skeletal muscle repair, three-dimensional piezoelectric fibrous scaffolds, often based on PLLA or PVDF, support myoblast differentiation and alignment. Experimental studies reveal that piezoelectric stimulation mimics the electromechanical cues of native muscle, enhancing myotube formation and contractility, with scaffolds enduring up to 40 % strain without loss of function, aiding volumetric muscle loss recovery. Regarding skin wound healing, biodegradable piezoelectric PLLA scaffolds activated by ultrasound promote fibroblast proliferation, collagen synthesis, and angiogenesis. In vivo findings indicate accelerated epithelialization and reduced scarring in rodent models, attributed to electrically induced anti-inflammatory responses and enhanced growth factor expression like VEGF.Fig. 8Illustration of the main effects of the application of piezoelectric materials in tissue engineering of bone, cartilage, nerves, skeletal and cardiac muscle and skin, and proposed future approaches for improving the field of research.Fig. 8
The characteristics of piezoelectric polymers have been extensively reviewed, revealing their significant potential in enhancing tissue healing processes through electrical stimulation. Biological structures, across various organizational levels (atomic, macromolecular, cellular, tissues, among others), commonly exhibit spiral morphologies. Since spirals inherently lack a center of symmetry, the piezoelectric phenomenon is an intrinsic property in many, if not all, biological systems, including bone, tendon, skin, cartilage, arteries, and ligaments, among others. Consequently, most biological tissues naturally experience electrical stimuli during physiological processes. Piezoelectric materials can effectively mimic the mechanical and electrical properties of native tissues.
In particular, under mechanical stress, bone generates electrical charges, which are known to promote essential processes such as cell alignment, migration, and differentiation, and are necessary for bone formation and healing. The piezoelectric property is also thought to play an important role in the control of cell nutrition, enzyme activation or suppression, and migration and proliferation of cells. This review highlights various piezoelectric polymers, such as PLA, PVDF, and others, that have been employed in tissue engineering research. However, the piezoelectric properties and their specific relationships to cell proliferation are often not reported. It is, therefore, imperative to establish clear correlations between piezoelectric coefficients such as d33 or d31 or other relevant parameters to cellular responses, particularly proliferation and differentiation.
Nevertheless, our comprehensive examination of the literature reveals that, despite considerable interest in piezoelectric biomaterials, there is a paucity of studies directly associating quantitative piezoelectric coefficients (d and g values) with cellular characteristics such as growth, proliferation, and differentiation. This absence is notable as it underscores a fundamental disjunction between the materials science community, which concentrates on characterizing and enhancing piezoelectric response, and the biological community, which aims to utilize bioelectrical signals to regulate cell fate. To move the field forward, it is important to close this gap.
Currently, scaffold and device design is predominantly empirical, depending on material type, processing conditions, or morphology, lacking a definitive predictive framework. This difference in piezoelectric coefficients that occurs when cells grow and differentiate is a significant challenge for designing biomaterials for regenerative medicine in a logical manner. This is especially true because researchers are increasingly investigating piezoelectric materials for tissue engineering and regenerative medicine. Their ability to convert mechanical forces into localized electrical signals could be utilized to regulate cell behavior. It is challenging to engineer materials that consistently provide the specific type, magnitude, and frequency of electrical signals known to influence processes such as stem cell lineage commitment, proliferation, and extracellular matrix production without this correlation.
Piezoelectric coefficients are highly sensitive to the organization of materials on different scales. To control biological responses, we need to understand how these dependencies function. Dipole orientation at the molecular level determines whether piezoelectric domains enhance or cancel each other out, thereby directly influencing charge density generation under stress. At the nanoscale, characteristics such as crystallite size, morphology, and domain boundary density influence dipole mobility and polarization stability, leading to localized fluctuations in piezoelectric output. At larger length scales, the long-range arrangement of dipoles in bulk material determines whether the electrical fields that cells experience are uniform or variable and unpredictable. Each of these structural factors changes the effective d and g values, which in turn control the strength, frequency, and spatial distribution of electrical signals at the cell-material interface.
For cells, electrical microenvironments serve as bio-instructive signals that control ion channel activity, intracellular signaling cascades, cytoskeletal organization, and ultimately gene expression. Electrical signals are known to influence the differentiation of mesenchymal stem cells. Small changes in the piezoelectric response could mean the difference between encouraging osteogenesis and adipogenesis, or between keeping the ability to grow and pushing differentiation. It is impossible to make scaffolds that give repeatable and optimized stimuli without a mechanistic framework that connects material properties (d and g values) to specific cell outcomes.
Finding this link has big effects. Hence, a deeper understanding of the relationship between piezoelectric properties and cell growth is critical. Materials with higher piezoelectric coefficients can generate stronger electric fields under the same level of mechanical stress, potentially providing more robust stimulation for cell types that require higher activation thresholds. However, it is important to note that beyond an optimal field strength, higher signals could have a detrimental effect on cell growth and proliferation. Further research is needed to determine the most effective piezoelectric coefficients and types of mechanical stimulation appropriate for different tissue types. By engineering scaffolds that replicate the natural piezoelectric characteristics of biological tissues, researchers can aim to provide the necessary electrical cues to guide cell behavior in a manner similar to native environments. Therefore, the piezoelectric coefficient is a key parameter in the design of tissue engineering applications. Selecting materials with appropriate piezoelectric properties can enhance cell signaling and promote tissue regeneration, positioning this as a promising area of research in regenerative medicine and bioengineering.
We can transition from trial-and-error fabrication to rational material design by clarifying how molecular orientation, nanoscale morphology, and long-range ordering influence the piezoelectric response. This would make it possible to make scaffolds and devices that send cells controlled, predictable, and adjustable electrical signals. This ensures that stimuli are consistent across batches, can be produced in large quantities, and are logically suited to specific therapeutic uses. This kind of framework would not only speed up discovery but also connect basic material science with clinical translation. This would enable next-generation implantable systems to utilize the body's own mechanical activity to regulate tissue regeneration.
This kind of framework would turn piezoelectric biomaterials from a novelty into a platform technology, connecting materials science and cell biology. The effects would be faster discovery, shorter development time, consistent results across batches, and trustworthy clinical translation. This research will ultimately establish the groundwork for next-generation implantable systems and intelligent scaffolds that utilize the body's intrinsic mechanical activity to facilitate tissue regeneration in a precise, individualized, and scalable fashion.
While the piezoelectric coefficient remains pivotal, the clinical implementation of piezoelectric implants hinges on more than just piezoelectric output or efficiency. Expedient secondary functions, such as stretchability to accommodate dynamic tissue movements and self-healing capabilities after cyclic mechanical loading and unloading, are crucial to ensure device reliability and prevent failure in vivo. For example, incorporating elastomeric components into piezoelectric composites can enhance fatigue resistance, allowing sustained performance in load-bearing applications like bone or muscle scaffolds. Additionally, the biodegradability of piezoelectric polymers profoundly affects tissue engineering outcomes. Biodegradable options like PLLA or PHBV degrade over time via hydrolysis, releasing non-toxic byproducts that integrate with regenerating tissue, eliminating the need for secondary surgeries. Controlled degradation rates, matched to tissue ingrowth, can provide temporary electrical stimulation while scaffolding degrades, fostering seamless tissue maturation. However, overly rapid degradation may compromise structural integrity and electrical output prematurely, potentially hindering regeneration, whereas slow degradation risks chronic inflammation. Thus, tuning molecular weight and copolymer ratios offers a pathway to optimize biodegradability for specific applications.
Emerging piezoelectric biomaterials, such as self-assembling peptides and natural biopolymers, present novel avenues with excellent biocompatibility. Piezoelectric peptides, derived from amino acids like glycine or diphenylalanine, form nanostructures with high d33 values and inherent bioactivity, promoting cell adhesion without external poling. Other biopolymers, including chitosan and cellulose derivatives, exhibit shear piezoelectricity and can be engineered into hydrogels or films for seamless integration with soft tissues, offering low immunogenicity and tunable degradation. These materials could revolutionize minimally invasive therapies by enabling injectable, self-powered systems.
Looking ahead, specific research directions in piezoelectricity and tissue engineering should focus on material design challenges, such as developing multimodal scaffolds that combine piezoelectricity with magnetic or photothermal stimuli for synergistic effects. Addressing scalability and standardization of poling processes to achieve consistent d33 values across batches is essential for clinical translation. In application technology, in vivo long-term studies are needed to evaluate immune responses and integration in dynamic environments. Personalized approaches, leveraging 3D printing with patient-specific piezoelectric gradients, could optimize outcomes for complex defects. Moreover, exploring hybrid systems with stem cell encapsulation and real-time monitoring via integrated sensors will bridge gaps in understanding electromechanical signaling. Overcoming biocompatibility hurdles for non-degradable components and ensuring regulatory compliance will accelerate adoption, ultimately transforming regenerative strategies into viable clinical therapies.NOMENCLATUREUNITScijklEFourth order elasticity tensor of stiffness constant at a constant electric fieldNm^−2^dijkThird order tensor of piezoelectric charge constantsCN^−1^ or mV^−1^DiElectric displacement componentCm^−2^eijkThird order tensor of piezoelectric constantsCm^−2^EiElectric field componentVm^−1^gijkPiezoelectric voltage constantm^2^C^−1^ or VmN^−1^hijkPiezoelectric constantVm^−1^ or NC^−1^sijklDElastic compliance at constant electric displacementm^2^N^−1^sijklEElastic compliance at constant electric fieldm^2^N^−1^SijComponents of second order Strain TensorβijτImpermittivity constants at constant stressVmC^−1^βijτImpermittivity constants at constant strainVmC^−1^τijComponents of second order Stress TensorNm^−2^εijComponents of Second order tensor of permittivity constantsCV^−1^m^−1^
Luís Martins: Writing – review & editing, Writing – original draft, Conceptualization. Ana Isabel Barbosa: Writing – review & editing, Writing – original draft, Funding acquisition, Conceptualization. Vítor Manuel Correlo: Writing – review & editing, Writing – original draft, Supervision, Conceptualization, Funding acquisition. Mrinal Bhattacharya: Writing – review & editing, Writing – original draft, Conceptualization. Rui Luís Reis: Writing – review & editing, Supervision, Resources.
This review article synthesizes existing literature on piezoelectric materials and their applications in tissue engineering. No original research involving human participants or animals was conducted for this study. Therefore, ethics approval and consent to participate were not required.
The authors declare the following financial interests/personal relationships which may be considered as potential competing Rui L. Reis is an associate editor for Bioactive Materials and had no involvement in the peer review of this article and had no access to information regarding its peer review. Full responsibility for the editorial process for this article was delegated to another journal editor. The remaining authors declare no conflict of interests.