Ultrasound-activated piezoelectric biomaterials for cartilage regeneration

Graphical abstract

Abstract

Due to the low density of chondrocytes and limited ability to repair damaged extracellular matrix (ECM) in cartilage, many patients with congenital or acquired craniofacial trauma require filler graft materials to support facial structure, restore function, improve self-confidence, and regain socialization. Ultrasound has the capacity to stimulate piezoelectric materials, converting mechanical energy into electrical signals that can regulate the metabolism, proliferation, and differentiation of chondrocytes. This unique property has sparked growing interest in using piezoelectric biomaterials in regenerative medicine. In this review, we first explain the principle behind ultrasound-activated piezoelectric materials and how they generate piezopotential. We then review studies demonstrating how this bioelectricity promotes chondrocyte regeneration, stimulates the secretion of key extracellular components and supports cartilage regeneration by activating relevant signaling pathways. Next, we discuss the properties, synthesis, and modification strategies of various piezoelectric biomaterials. We further discuss recent progresses in the development of ultrasound-activated piezoelectric biomaterials specifically designed for cartilage regeneration. Lastly, we discuss future research challenges facing this technology, ultrasound-activated piezoelectric materials for cartilage regeneration engineering. While the technology holds great promise, certain obstacles remain, including issues related to material stability, precise control over ultrasound parameters, and the integration of these systems into clinical settings. The combination of ultrasound-activated piezoelectric technology with other emerging fields, such as Artificial Intelligence (AI) and cartilage organoid chips, may open new frontiers in regenerative medicine. We hope this review encourages further exploration of ultrasound-activated strategies for piezoelectric materials and their future applications in regenerative medicines.

1. Introduction

Electrical stimulation (ES) is widely used method for in vitro and in vivo biomedical applications, exerting beneficial effects on cellular functions such as metabolism, proliferation, and differentiation [1]. For instance, cardiomyocytes contract in response to electrical stimulation, nerve cells regulate muscle activity and convey messages via electrical currents, and the piezoelectric effect in bone and cartilage tissues facilitates structural changes in the ECM and cellular rearrangement under pressure stimulation. However, a major challenge in cartilage regeneration is the lack of standardized ES activation patterns that can efficiently promote chondrocyte differentiation and cartilage regeneration [2]. Due to the deep location of cartilage within the tissue and the significant individual variation in joint and skin thickness, the penetration capability of exogenous ES remains uncertain. Additionally, short-term invasive ES is not well-suited for supporting the prolonged growth cycles of chondrocytes and poses an increased risk of infection. Thus, it is urgent to develop new, non-invasive ES activation pathways to promote cartilage regeneration efficiently.

Ultrasound-driven microcurrents using piezoelectric biomaterials is emerging as a promising approach for cartilage regeneration. A pioneering work by Liu et al. demonstrated successful cartilage regeneration in rabbits through the utilization of limb movement to generate piezoelectric stimulation in joint defects [3]. However, not all cartilage types are involved in joint or limb movement. Cartilage in areas such as the ear, nose, and ribs primarily provide structural support, and thus cannot rely on limb movement to generate sufficient bioelectricity from grafted piezoelectric biomaterial for promoting cartilage repair and regeneration. To address this, Lu et al. introduced a radio-stimulation method employing ultrasonic vibrations to trigger piezoelectric nanomaterials, thereby facilitating the expansion and maintenance of neural stem cells. Genchi et al. demonstrated the potential of ultrasound in promoting the differentiation of human SaOS-2 osteoblasts [4,5]. These studies showcase the feasibility of using ultrasound-activated piezoelectric biomaterials for endogenous power generation in cartilage regeneration engineering, providing a theoretical basis for chondrocyte regeneration across various anatomical sites.

Despite the advantages of electrical signal modulation and the structural-mechanical properties of piezoelectric materials, these materials often elicit adverse chronic foreign body reactions at the graft site, leading to fibrous formation that may compromise the therapeutic outcomes [6,7]. Consequently, there exists an urgent necessity to comprehend the mechanisms underlying ultrasound stimulation of cartilage regeneration via piezoelectric materials and to identify piezoelectric materials possessing suitable structure and biocompatibility for remotely modulating cartilage regeneration.

This review begins by elucidating the principles behind ultrasound-activated piezoelectric materials for cartilage regeneration, particularly their role in enhancing cartilage cell proliferation. It then categorizes the various piezoelectric materials used in cartilage regeneration, evaluating the advantages and limitations of inorganic, organic, and composite types. Finally, the review summarizes recent advancements in the application of ultrasound-activated piezoelectric biomaterials for cartilage tissue regeneration, highlighting their potential and identifying critical research gaps. The arm is to provide researchers with valuable insights for material selection and strategy development in bioelectricity-related biomedical applications.

2. Mechanisms

2.1. Non-centrosymmetric crystal structure causes material piezoelectricity

The piezoelectric effect is a characteristic property of most non-centrosymmetric crystals. However, not all non-centrosymmetric crystals exhibit significant piezoelectric effects. The manifestation of piezoelectricity depends on the specific structure of the crystal. Factors such as lattice structure, atomic arrangement, and stress distribution significantly influence the piezoelectric response. In recent years, the study of the relationship between crystal structure and piezoelectricity has led to significant progress in material design and property modulation. Great efforts have been devoted to introducing phase instabilities for enhancing piezoelectricity by facilitating polarization variation, including the establishment of morphotropic phase boundary (MPB) and local structure heterogeneity (LSH) or polar nanoregions (PNR) [8] (Fig. 1).

Fig. 1.

Strategies for modulating the piezoelectricity of crystal structures. (A) High-angle annular dark field (HAADF) images of pristine, Fe-doped, and Mn-doped KTN crystal microstructures differentiated by the MPB strategy tuning. a-c) The small bright dots and the large bright spots represent the A-site and B-site ions, respectively, and yellow arrows mark the mapping of the polar vectors; c) the projected polarization vectors in the Mn-doped samples are highly disordered, whereas the b) Fe-doped samples maintain almost uniform polarization orientation; d-f) statistical analyses of the polarization distributions of the d) pristine; e) Fe-doped, and f) Mn-doped samples, which indicate the inhomogeneity in the crystal structure of the f) Mn-doped samples. (reproduced with permission from Ref.[9] Copyright 2024 Nature); (B) a) schematic of the LSH strategy for synthesizing 2D bismuth-based chalcogenide nanosheets; b) phase diagram of MA₃Bi₂Cl₉-PEG (MBCP); c) phase hysteresis line and amplitude butterfly curve of MBCP; d) piezoresponse amplitude of MBCP (reproduced with permission from Ref.[10] Copyright 2025 ELSEVIER); (C) observation of polar nanoregions (PNR) in implanted PbZrO₃ (PZO); a) polarization vector mapping based on helium ion injection into PZO at a dose of 2.5 × 10¹⁵ ion/cm²; the corresponding out-of-plane b) amplitude and c) phase PFM images of the injected PZO membrane; d) schematic of an injected PZO cell with a tetragonal phase and e) schematic of a PZO cell with an antiferroelectric structure (reproduced with permission from Ref.[11] Copyright 2023 Applied Physics Reviews). (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

MPB refers to the coexistence of different crystalline phases (e.g., rhombic, tetragonal, monoclinic, etc.), and the polarization vectors between different phases achieve rotation at lower energy under external electric field or stress, thus greatly enhancing the piezoelectric response. Constructing MPBs by adjusting the ratio of transition metals with the same lattice occupancy is an important method to enhance piezoelectricity. Tan et al. doped potassium tantalum niobate (KNN) single crystals with transition metals (Fe, Mn) and found that doping-induced lattice distortions and changes in the local charge distribution significantly enhance the piezoelectric coefficients [9]. Abramov et al. used SrTiO₃ doped BiFeO₃-BaTiO₃ to improve the electromechanical response of the material while maintaining the quasi-cubic structure of the crystals by dissolving polar rhombic and tetrahedral nanoregions in a nonpolar matrix [12]. Lin et al. induced the O-T phase boundary and enhanced the relaxation properties by introducing Ba(Ti_0.7Sn_0.3)O₃ and NaSbO₃ to form a localized heterostructure, which enhanced the piezoelectricity coefficient, d₃₃, to 423 pC/N [13].

LSH enhances the piezoelectricity by introducing localized inhomogeneities in the lattice distortions caused by the chemical composition. For example, Zhu et al. effectively disrupted the centrosymmetry of MA₃Bi₂Cl₉ to form a non-centrosymmetric crystal structure through the introduction of methylamine cation. This structural modification significantly enhances the piezoelectric properties, enabling a more robust charge separation effect under ultrasonic excitation [10].

PNRs are dynamically existing nanoscale polar regions in relaxing ferroelectrics, whose polar orientations can be rapidly reoriented under an applied electric field or stress. This dynamic response lowers the energy barrier for polarization flip and makes the material more susceptible to polarization, thus enhancing the piezoelectric sensitivity. Luo et al. induced an ordered-disordered phase transition in the antiferroelectric material PbZrO₃ to form a PNR by helium ion injection, which not only lowered the domain-flip energy barriers, but also increased the polarization strength by enhancing the tetragonal distortion, and enhanced the energy storage density from 20.5 J/cm³ to 62.3 J /cm³[11].

It is worth mentioning that MAGUS 2.0, a crystal structure prediction software developed by Han et al. can combine symmetry principle and machine learning to extract global and local features of the explored crystal structures using group theory and graph theory, which significantly improves the prediction efficiency of complex crystal structures (e.g.,γ-boron and violet phosphorus) [14]. Future studies are still needed to combine in situ characterization and computational simulation to further reveal the coupling mechanism of MPB, LSH and PNR, in order to promote the application of bio-piezoelectric materials in biomedical fields.

2.2. Mechanisms of electrical stimulation for cartilage repair

The cartilage ECM is composed primarily of collagen, which provides structural support, and proteoglycans, which are known for their high-water absorption capacity. Among these, type II collagen, is recognized as a natural piezoelectric material and is often used as a key indicator for assessing cartilage regeneration. The piezoelectric effect in cartilage is attributed to the non-centrosymmetric structure of type II collagen, which allows it to respond to mechanical loads, thereby affecting cell membrane receptors and ultimately the cell nucleus [15].

When stress is applied to piezoelectric materials, an electrical potential is generated across their two opposing surface [16]. To precisely control the application and intensity of stress while minimizing invasiveness, ultrasound can be employed as an alternating source of mechanical force [17]. Ultrasound can penetrate deeply into tissues and generate acoustic pressure that induces charge migration within the piezoelectric crystals. This results in charge accumulation on the opposing surfaces, thereby activating cellular signaling pathways. These pathways, in turn, promote various cellular behaviors such as growth, proliferation, differentiation, and secretion [18]. At the tissue level, piezoelectric materials, such as films or scaffolds, can be stimulated by ultrasound beneath the skin and muscle tissue, which generates microcurrents on the surface of the grafted cartilage tissue encapsulated in piezoelectric material. This microcurrent mimics the electrophysiological microenvironment of cartilage tissue, regulates the proliferation and maturation of cartilage membranes and chondrogenic progenitor cells, and promotes chondrocyte proliferation and facilitates the synthesis of ECM molecules such as glycosaminoglycans, hyaluronic acid, and COL2 [[19], [20], [21]]. This method fosters cartilage tissue regeneration by generating bioelectrical stimulation via controlled ultrasound energy, eliminating the need for an external power source [22]. For example, culturing MC3T3-E1 preosteoblasts under mechanical strain of 110 ppm and electrical voltage of 0.115 mV significantly enhanced cell proliferation by up to 25 % [23]. It is noteworthy that both pre-osteoblasts and chondrocytes are differentiated from osteoprogenitor cells. In the strategy of ultrasound activation of piezoelectric biomaterials to generate microcurrents, it is important to emphasize that low-frequency US benefits chondrocytes, whereas high-frequency US tends to induce chondrocyte apoptosis. Ricotti et al. found that 1 MHz US was more effective than 5 MHz US in promoting the expression of cartilage-forming genes such as COL2A1, aggrecan (AGG) and SOX9 [24].

ES can modulate chondrocyte behavior through several mechanisms, including mechanical activation of stretch-activated calcium channels (SACC, also known as Piezo channels), voltage-gated Ca²⁺ channels (VGCC), and the activation of transforming growth factor β (TGF-β) and Wnt/β-catenin signaling pathways (Fig. 2), etc [25,26]. Liu et al. showed that mechanical stimulation of piezoelectric scaffolds activated VGCC and SACC channels, leading to the generation of electrical signals. These signals, on one hand, trigger the release of large amounts of Ca²⁺ into the cytoplasm through the cell membrane and endoplasmic reticulum (ER), leading to a transient elevation of intracellular calcium [27]. Electrical stimulation also dephosphorylates phosphorylated nuclear factor (NF-AT) through activation of the calmodulin (CaM) and calcineurin (CaN) pathways, translocating it to the nucleus. This promotes gene transcription and the production of TGF-β and bone morphogenetic protein type 2 (BMP-2). The produced TGF-β plays a critical role in promoting cell proliferation and differentiation, ECM synthesis, and inflammatory repair processes. It also affects bone and cartilage formation by inhibiting drosophila mother's resistance to decapentaplegic (SMAD) protein expression and suppressing chondrocyte fibrosis [28]. Inhibition of TGF-β-induced Wnt/β-catenin signaling has been shown to enhance the chondrogenic capacity of plasma hydrogel mesenchymal stem cells (MSCs) [29]. The dishevelled proteins (DVL) in the cytoplasm receive upstream signals that stabilize β-Catenin protein by inhibiting the function of a protein complex comprising activator protein C (APC), Axin (a scaffolding protein responsible for axis inhibition), and glycogen kinase 3β (GSK3β). This allows itself to enter the nucleus and bind to the β-Catenin gene, which initiates the transcription of downstream target genes [30]. This pathway ultimately leads to an increase in the secretion of collagenase type II (COL2), AGG, SRY-Box transcription factor (SOX) and glycosaminoglycan (GAG), while decreasing the secretion of collagenase type I (COL1) and production of matrix metalloproteinase (MMP) [31].

Fig. 2.

Diagram of the molecular mechanisms by which ultrasound-stimulated piezoelectric materials remotely modulate chondrocyte regeneration.

ES can also directly affect intracellular membrane potential, resulting in increased intracellular reactive oxygen species (ROS) levels [32]. To maintain ROS levels suitable for cell growth, excess ROS are detected in the cytoplasm by the nuclear factor E2-related factor 2 (NRF2) and the two kelch-like ECH-associated protein 1 (KEAP1) complex, which triggers the release of NRF2 from KEAP1, allowing NRF2 to regulate the expression of antioxidant-related substances such as guanosine triphosphate cyclase (GCH1), sulfoximine (Srx), thioredoxin reductase (TrxR), peroxiredoxin-1 (Prx-1), catalase, superoxide dismutase (SOD), glutathione peroxidase (Gpx), heme oxygenase 1 (HO-1), and nicotinamide adenine dinucleotide phosphate (NADPH) [33]. These antioxidant substances alleviate oxidative damage to the cells to a certain extent, for instance, GCH1-mediated synthesis of 5,6,7,8-tetrahydrobiopterin (BH4) significantly reduces ROS production [34]. However, once this balance is disrupted, such as during an imbalance in intracellular ROS production, prolonged pulsatile stimulation can lead to excessive ROS production in the mitochondria. This causes the accumulation of dysfunctional mitochondria in chondrocytes, resulting in the release of cytochrome C (Cyt C), which activates the caspase signaling pathway and ultimately leads to apoptosis [35,36]. Although ES can promote local vasodilation, enhance vascular permeability, increase blood flow to tissues, and provide nutrients to transplanted cartilage, prolonged or high-frequency US exposure may transmit excessive energy, potentially harming cartilage regeneration [37].

2.3. Ultrasound-mediated ES device, parameters and piezoelectricity test instrument

Ultrasound-mediated electrical stimulation typically requires specialized equipment to achieve precise control [38] (Fig. 3). For example, Chang et al. designed a programmable electrically stimulated cell culture device for cartilage cells culture [27], while Chen et al. independently developed a six-channel ultrasound stimulation system with low-intensity pulsed ultrasound [29]. Rong et al. performed remote reception experiments of piezoelectric materials in a tank filled with deionised water to characterise the US reception performance of the AlN piezoelectric micromachined ultrasonic transducers array, and the results show that the reception sensitivity of this piezoelectric material underwater is about 1 V/MPa [39].

Fig. 3.

Representative diagram of ultrasound-mediated ES devices and techniques. (A) ES devices for chondrocyte culture (reproduced with permission from Ref.[27] Copyright 2021 ELSEVIER); (B) a) the six-channel ultrasound stimulation system used for in vitro LIPUS (low intensity pulsed US stimulation) stimulation; b) the parameters of in vitro LIPUS stimulation: Frequency = 1 MHz, PRF (pulse repetition frequency) = 0.33 Hz, Duty cycle = 20 %, with sound intensity ranging from 30 to 250 mW/cm² for 20 min (reproduced with permission from Ref.[29] Copyright 2023 BMC); (C) experimental setup for evaluating AlN piezoelectric micromachined ultrasonic transducers array as wireless power receiver (reproduced with permission from Ref.[39] Copyright 2022 Nature).

Regarding ultrasound frequency, high-frequency ultrasound in the range of 1–3 MHz typically used in diagnostic imaging may cause thermal damage to cartilage tissue during prolonged exposure [40]. In general, ultrasound intensity is recommended to be kept below 0.5 W/cm² [29,41] to avoid damage. Lower frequencies, such as 40 kHz, are safer and more suitable for penetrating joint tissues and activating piezoelectric materials. This is because the relaxation time of human tissues is about 0.1 ∼ 1 μs, and low-frequency US can effectively reduce the absorption of heat by tissues [42]. Although direct stimulation is more convenient, it is usually limited by imprecise targeting and the potential for tissue damage from uneven distribution. Farooqi et al. demonstrated through mathematical modeling that piezoelectric hydrogels help distribute the internal electric field more evenly [37]. However, most piezoelectric biomaterials degrade over time, and their piezoelectric properties decrease, reinforcing the need for dynamical adjustment of US stimulation parameters rather than fixed setting [43].

Notably, researchers typically use electrochemical workstations to assess the piezoelectricity of materials. This involves applying force to a material block and recording the resulting charge using a charge amplifier to measure piezoelectric coefficients. The reverse piezoelectric effect is primarily evaluated using laser displacement sensors. Piezoresponse force microscopy (PFM) has become a powerful technique for analyzing localized piezoelectric effects at micron and nanometer scales. In PFM, a conductive tip scans the surface of a material while an alternating current (AC) voltage is applied, measuring the mechanical deformation or electrical response. As a high-resolution scanning probe technique, PFM provides detailed understanding of piezoelectric coefficients at the microscopic level, making it particularly effective in the study of small or heterogeneous materials [44].

3. Piezoelectric biomaterials

Piezoelectric biomaterials suitable for cartilage tissue engineering can be typically categorized into three types based on their molecular composition and degradable nature: inorganic piezoelectric materials, organic piezoelectric materials, and piezoelectric composite, such as piezoelectric hydrogels (Table 1). Most inorganic piezoelectric materials are non-degradable but have high piezoelectric coefficients [45]. Organic piezoelectric materials, on the other hand, offer superior biocompatibility, but they tend to have lower piezoelectric coefficients compared to inorganic piezoelectric materials, limiting their ability to generate sufficient piezoelectrical charges. Piezoelectric composites, such as composite hydrogels, combine the flexibility of organic materials with the enhanced piezoelectric properties provided by the incorporation of inorganic nanoparticles with high piezoelectric coefficients.

Table 1.

Piezoelectric biomaterials and their applications in cartilage regeneration.

Types	Materials	Morphology Design	Dimension	Synthesis Route	Working Condition	Piezoelectric Coefficient	Ref.
Organic piezoelectric materials	γ-glycine	PVA-glycine-PVA sandwich heterostructure	thickness: ≈ 30 mm	direct solidification from the mixture solution at 60°C	The Vpp output retained a constant value of ∼ 4.1 V when subjected to more than 10,000 cycles of 30 N impulse force application	d33: 5.3 pC/N	[46]
	Diphenylalanine (FF)	A vertical FF microrods arrays	1.25*1.25 cm2 gold-coated silicon substrate	By controlling water diffusion during the seed film formation, seeds with a majority of vertical domains with antiparallel polarizations could be achieved, and the application of an electric field made the polarizations uniform	Under an applied force F¼ 60 N, the output open-circuit voltage (Voc) and short-circuit current (Isc) reached 1.4 V and 39.2nA	d33: 17.9 pm/V	[47]
	Fmoc-FF	nanotubes and film	diameters: ≈ 100 mm, height: ≈ 667 ± 170 nm	a solvent-based method	A loading force of 10 nN at specific locations and a 12.5 kHz (Fmoc-FF sample) AC voltage was swept from 0 to 60 V/ms in 6 V/ms increments while recording LPFM response.	d15: 33.7 pm/V	[48]
	Collagen	rat tail tendon film	diameters: ≈ 184 nm, thickness: ≈10 μm	tendon was harvested directly from the tail of a 4 weeks old with embedding in epoxy	conductive cantilevers having a nominal spring constant and resonant frequency of 2.8 N/m and 75 kHz	d15: 6.21 pm/V	[49]
	Elastin	aortic elastin	thickness: ≈ 0.62 mm	a purified elastic fiber network was obtained from a porcine thoracic aorta	A 5 V AC voltage near resonance	deff: ∼ 1 pm/V	[50]
	M13 phage	4E engineered M13 phage monolayer	a long rod-like shape, 880 nm in length and 6.6 nm in diameter	phage can be produced simply and economically by infecting bacteria. Thin film preparation by dropcasting	The peak current increased linearly for the displacement rate of the strain force increases from 0.06 to 0.1, or the strain rate from 0 to 0.4 s−1	d33: 11.2 pm/V, deff ≈ 3.9 pm/V (for films greater than 100 nm thick)	[51]
	Cellulose	Vertically aligned CNC film	a 12 × 5 mm2 CNC film	to uniaxially align CNCs effectively, an ethanol −water binary solvent was used to suppress the formation of a chiral structure and used for CNC alignment	5 kV DC voltage	d33: 19.3 pm/V	[52]
	Chitin	β-rich chitin film	thickness: ≈ 35 μm, 6 × 10 cm2 film	preparation of chitin self-supporting films by centrifugal heating after solubilization	at different frequencies (0.2 – 4.1 Hz) and a fixed pressure of 19.8 kPa, the output current density monotonically increased with vibration frequency up to 4.1 Hz	deff: 3.986 pm/V	[53]
	Chitosan	Neutralized chitosan film	thickness: ≈ 15 μm	solvent-casting technique	excited at different AC voltages (from 0.5 to 5 V)	d33: 15.56 pC/N	[54]
	PVDF (poly(vinylidene fluoride))	Film	diameters: ≈ 12 mm, height: ≈ 24 mm	3D technique	24 h of cocultivation	—	[55]
	P(VDF-TrFE)	Nanofiber	diameter: ≈ 1.24 ± 0.13 µm	electrospinning	12 d of cocultivation	d33: 11.1 pm/V	[56]
	PHBV (poly- 3- hydroxybutyrate-3-hydroxy valerate)	—	a pore diameter: 1.139 μm	electrospinning	co-culture with MSC-differentiated chondrocytes	d33: 0.4 pC/N	[57]
	Poly-l-lactic acid (PLLA)	Fibers	diameters: ≈ 860 ± 110 nm	electrospinning	co-culture with hFob cells	—	[58]
Inorganic piezoelectric materials	PZT	Film	thickness: ≈ 100 µm	hydrothermal method	< 5 Hz bending frequency	16 V maximum output voltage	[59]
	(BaTiO3) BT	NPs	diameter: 58 ± 15 nm	solvothermal process and thermal annealing	1 MHz, 1 W/cm2 ultrasonic wave	—	[60]
	ZnO	NPs	average diameter: 30 – 80 nm, thickness: 0.1 and 20 µm	hydrothermal method	18 weeks in vivo	15.1 to 51.9 pm/V	[61]
	KNNSe	ceramic	thickness: ≈ 1 mm, diameter: ≈ 10 mm	solid phase sintering	3 d of cocultivation	d33: 120 pC/N	[62]
	BiFeO3 (BFO)	single crystals	0.3 × 0.2 × 0.02 mm3	spontaneous crystallization	20°C / 293 K	d33: 100  pm/V	[63]
	NbO3	ceramic	density: 4.25 g/cm3	air sintering	air-fired	d33 = 80 pC/N	[64]
	SiC	nanoporous film	the average diameter of nanohole: 73.67 nm	anodic oxidation	apply 1 MPa pressure vertically	—	[65]
	Aluminium Nitride (AIN)	films	thickness: ≈ 30 nm	deposition	beam epitaxy	d33 = 5.1 pm/V	[66]
	GaN	nanowires	the average length: 5 µm	vapor–liquid– solid (VLS) growth	2 to 12 Hz	—	[66]
	Al/Ga-doped h-BN	bulk	thickness: ≈ 50–––120 nm	doping approach	—	d33 = − 24.214 pC/N	[67]
	CdS	NPs	the lattice stripe: 0.213 and 0.344 nm	deposition	40 kHz, 120 W (ultrasonic)	—	[68]
	CdSe	thin-film chip	diameter: 13 mm	magnetron sputtering	light wavelength:1064 nm	d33 = 76 pm/V	[69]
			diameter of CdSe crystallite: 3–––5 nm
Piezoelectric composites	Cs/Gel/PHA/PBT	hydrogels	thickness: ≈ 300 μm	mixing method	10 Hz and 1 kPa pressure	Output voltage: 0.2–––0.8 V	[70]
	Alg-DA/Ac-β-CD/gelatin	hydrogels	can be injected by a 23 G needle	mix-lyophilize-mix	10T/s, 15 Hz (pulsed electromagnetic fields)	—	[31]
	(NdFeB)/Poly-GelMA-HAMA	hydrogels film	various kinds of shapes	mixing method	20 % deformation, 1 Hz (mechanical stimulation)	—	[71]
	CoFe2O4/Methacrylated Gellan Gum (GGMA)/poly(vinylidene fluoride) (PVDF)	microspheres hydrogels	thickness: 1 μ m	electrospray and mixing method	—	d33 = 22 pC/N	[72]
	Silk Protein Hydrogel	nanocomposite film hydrogels	thickness: 0.81 mm	hydrothermal process and casting method	1 Hz and 15 N pressure	d33 = 320 pC/N (ZnONRs-silk)	[73]
			5 mm × 5 mm square structure
	ShortNanofibers of PLLA (NF-sPLLA) / Collagen Matrix	injectable hydrogels	can be injected by a 29 G needle	electrospinning and mixing	40 kHz, 0.33 w/cm2 (ultrasonic)	—	[74]
	BaTiO3/Gelatin Methacryloyl (GelMA)/Tilapia Fish Gelatin Hydrogel	hydrogels	diameter: 6 mm, height:10 mm	solvothermal reaction, thermal calcination process, mixing and light-curing	maximum compression strain: 80 %	32.7 mV at 0.5 Hz, 113 mV at 2.5 Hz	[75]
	BTNPs/GO Nanoflakes/VitroGel-RGD	injectable hydrogels	can be injected by a 22 G needle	hydrothermal process and Coating	1 MHz, 250 mW/cm2, 20 % duty cycle, 5 min	d33 = 118.6 pm/V	[24]

When ultrasound exposure is utilized for cartilage tissue regeneration, several critical factors should be carefully considered, especially compared to other tissues like nerves or bone: i) Ultrasound penetration and frequency: Cartilage has lower density and different acoustic impedance compared to bone, which is denser and more rigid. In contrast, nerve tissue, being softer, more elastic and linear in structure, may interfere with the efficacy of ultrasound-mediated piezoelectric therapy. ii) Biological response to ultrasound: Cartilage regeneration involves chondrocyte activity and ECM synthesis. Ultrasound parameters must be optimized to promote these processes while avoiding netative effects such as cellular damage or ECM degradation. These specific parameters used for cartilage may not be directly applicable to neural tissue, due to their distinct cellular composition and function. iii) Thermal effects: It is essential to ensure that the ultrasound parameters do not cause excessive heating, which could impede cartilage regeneration or cause tissue damage. While bone tissue can tolerate higher temperatures, cartilage requires more precise temperature control to prevent thermal damage. (iv) US duration and intensity: The optimal duration and intensity of ultrasound exposure for effective cartilage regeneration must be carefully investigated, taking into account the unique cell cycles and tissue-specific response. This is particularly crucial in differentiating between cartilage and other tissue types. (v) Monitoring and assessment: Imaging and histological analysis techniques for cartilage, bone and nerves are essential. Selection of the appropriate ultrasound pattern is critical to ensuring adequate tissue penetration and effective stimulation of cartilage, while minimizing damage to surrounding tissues.

3.1. Inorganic piezoelectric materials

Inorganic piezoelectric materials typically exhibit high piezoelectric coefficients and excellent mechanical properties, which make them compatible with bone and cartilage tissues. However, piezoelectric ceramic are inherently brittle and prone to fracture, and may contain hazardous, non-biodegradable substances such as lead or hydrogen fluoride [76]. Additionally, their piezoelectricity tends to decrease at the nanoscale, particularly when their sizes drops below 100 nm [77]. Zhou et al. observed that larger pores in piezoelectric fiber films can enhance their piezoelectric properties [65], although the brittleness of ceramic materials renders them vulnerable to damage, thereby shortening their service life. Although these materials are not ideal for modeling cartilage due to their lack of flexibility, their excellent piezoelectric properties and high stiffness make them particularly suitable for modeling hard tissues such as bone [7,78,79]. Another class of inorganic piezoelectric materials includes hexagonal wurtzite crystals, including ZnO, SiC, AlN, GaN, InN, BN, CdS, and CdSe [80,81] (Table 1).

To mitigate the risk of biotoxicity from degradation byproducts, lead-free inorganic piezoelectric materials, such as titanate (BaTiO₃), potassium sodium titanate (KNN), and zinc oxide (ZnO), are often preferred for cartilage regeneration scaffolds. However, the biosafety and potential degradation byproducts of these materials have not been fully elucidated [82]. As a tetragonal crystal with an asymmetric structure, BaTiO₃ has shown high compatibility with biological systems. It exhibits strong polar and piezoelectric effects under mechanical stress, with a piezoelectric coefficient up to 340 pC/N [83,84]. Chen et al. synthesized octahedral BaTiO₃ using a solvent-thermal growth method, resulting in even stronger piezoelectric effects. In a low-intensity pulsed ultrasound (LIPUS)-stimulated subcutaneous implantation model, Wu et al. utilized BaTiO₃-coated Ti₆Al₄V scaffolds to electrically stimulate macrophages. This approach promoted anti-inflammatory polarization and facilitated bone repair by inhibiting MAPK/JNK (c-JunN-terminalkinases, JNK) signaling and activating oxidative phosphorylation (OXPHOS) [85]. The incorporation of BaTiO₃ into composites is a common strategy to enhance the piezoelectric performance, extend degradation time, and improve their suitability for a range of biomedical applications.

3.2. Organic piezoelectric materials

Organic piezoelectric materials typically offer advantages such as high flexibility, low toxicity, ease of biodegradation, and the ability to form porous structures. However, these materials generally exhibit poor mechanical properties compared to inorganic counterparts, and their synthesis often requires harmful solvents. Among organic piezoelectric materials, polyvinylidene fluoride (PVDF) stand out as a flexible, non-toxic option suitable for cartilage tissue engineering [86]. However, PVDF is non-biodegradable, posing a risk of fibrous capsule formation around implants [87,88]. Other biodegradable organic piezoelectric materials include poly (L-lactic acid) (PLLA), polyhydroxybutyrate (PHB), cellulose, chitosan, amino acids and peptides (Table 1). These materials exhibit high biocompatibility and are less prone to fibrous cyst formation.

3.2.1. PLLA

Poly-L-lactic acid (PLLA) is biodegradable and biocompatible, possessing a piezoelectric shear coefficient (d₁₄) of 10 pC/N [89]. In 1995, Fukada et al. first demonstrated PLLA’s piezoelectric effect upon implantation [90]. PLLA exhibits strong mechanical properties and is used in orthopaedic applications, such as screws, where its degradation by-products are non-toxic and water-soluble [91]. Electrospinning technology can be used to produce PLLA fibers, generating micron-diameter fibers that are instrumental for creating fibrous scaffolds for the Organ on Chip with Enhancement (OOCE) strategy [92]. The piezoelectric properties of electrospun PLLA materials can be modulated by controlling factors such as the nozzle aperture, drum speed, voltage and solvent system. Various post-treatment methods can further modify the scaffolds to enhance pro-tissue regeneration, promoting cell adhesion and growth [93,94].

3.2.2. PHBV

Poly (3-hydroxybutyric acid-co-3-hydroxyvaleric acid) (PHBV), a member of the PHB family, exhibits high biodegradability [78]. The piezoelectric coefficient of PHBV is approximately 1.3 pC/N, comparable to that of human bone [95]. Although PHBV has a longer degradation period than other biocompatible polymers, its low piezoelectric coefficient and inadequate mechanical properties constrain its application in piezoelectric scaffolds [78]. Despite these limitations, PHBV substrates have demonstrated excellent biocompatibility and extended biodegradation periods, making them suitable as a matrix for cartilage and bone regeneration [96]. A significant challenge in employing PHBV for these purposes lies is its hydrophobic nature, which can impair cell adhesion [97].

3.2.3. Cellulose

Cellulose, a homopolymer of glucose, possesses a shear piezoelectric coefficient (d₁₄) of 0.2 pC/N [98] and a shear piezoelectric charge coefficient (d₃₁) of 0.1 pC/N [89]. Despite its high-water content, cellulose maintains good mechanical properties and is biocompatible, elastic and biodegradable. The porosity of cellulose can be enhanced using paraffin microsphere porogens, which creates an interconnected pore network that improves cell adhesion [99]. Additionally, the nanofibrous structure of cellulose can be modified by incorporating suitable porous agents. Methylcellulose, a derivative of cellulose and an injectable biopolymer, mimics the electrophysiological microenvironment of tissues, making it particularly suitable for cartilage tissue regeneration [97]. Methylcellulose strongly promotes cell adhesion, though it has a low piezoelectric coefficient. To address this, some researchers have blended methylcellulose with stronger piezoelectric materials to create composite materials that combine the benefits of both [100].

3.2.4. Chitosan

Chitosan, a biodegradable polysaccharide derived from deacetylated chitin, is extensively utilized in cartilage tissue regeneration and repair engineering due to its excellent biocompatibility, antimicrobial properties, mechanical strength, and low immune rejection. It is abundantly available, cost-effective, and easily biodegradable. Chitosan can be processed into various composites, which further enhance its versatility in tissue engineering applications [101,102]. Studies have illustrated that the chitosan degradation products have anti-apoptotic effects on chondrocytes, significantly decreasing pro-inflammatory cytokine expression and promoting the maintenance of cartilage morphology by supporting collagen II expression in chondrocytes [103,104]. One study by Gossla et al. combined alginate hydrogel with chitosan-based implantable scaffolds, creating scaffolds with an anisotropic, porous structure and good elasticity. This combination promoted chondrocyte attachment, chondrogenic differentiation, and ECM formation [105].

3.2.5. Amino acids (AAs)

Amino acids, the fundamental building blocks of peptides and proteins, exhibit piezoelectric properties due to the non-centrosymmetric nature of their crystal structures. Among the 20 naturally occurring AAs, experimental evidence supports piezoelectric effects in 17, with the eceptions being glutamine, phenylalanine, and tryptophan [106,107]. Glycine, in particular, has emerged as the most extensively studied piezoelectric AA. Its ability to undergo a phase transformation from a non-piezoelectric to a piezoelectric phase under an applied electric field is attributed to structural rearrangement in its crystal lattice and realignment of molecular dipoles. Glycine crystallizes into three distinct crystal structures, α-phase, β-phase, and γ-phase, depending on the crystallization conditions [108]. Specifically, both β-glycine and γ-glycine possess non-centrosymmetric crystal structures and demonstrate significant piezoelectric properties, whereas α-glycine crystal lacks piezoelectricity due to tis centrosymmetric crystal symmetry [109,110]. The piezoelectric charge coefficient of β-glycine is approximately 10 pm/V, comparable to that of traditional organic piezoelectric materials. Remarkably, the maximum piezoelectric coefficient of glycine surpasses that of widely used piezoelectric polymers, such as collagen and PVDF, by several folds [109]. A notable feature of β-glycine is its ability to maintain piezoelectric sensitivity even under ultra-low mechanical pressures [111].

The piezoelectric coefficients of AA-based materials span a broad range, from 0.5 pC/N to 178 pC/N, depending on their molecular and crystalline structures [112]. Zhang et al. proposed a self-assembly strategy to customize piezoelectric biomaterial films, achieving a piezoelectric voltage coefficient (g₃₃) of 252 × 10^-3 Vm/N for β-glycine films [113]. However, the practical application of glycine-based piezoelectric materials is limited by their inherent brittleness and rapid degradation in physiological environments [114]. In 2020, Hosseini et al. addressed this challenge by stabilizing β-glycine crystals through self-assembling glycine in an aqueous chitosan matrix using a straightforward solvent-casting approach. This approach efficiently prevented the spontaneous conversion of metastable β-glycine into the more stable α- or γ-phase, while maintaining a low loss factor of 0.18 [115]. Additionally, the application of an electric field also induces polarization, transforming the non-piezoelectric phase into a piezoelectric phase [116].

3.2.6. Peptides

Peptides, short chains of amino acids, also exhibit piezoelectricity due to their asymmetric crystal structures [117,118]. Notable examples of piezoelectric peptides include diphenylalanine (FF) [119], fluorenyl methoxycarbonyl diphenylalanine (Fmoc-FF) [48], cyclo-glycine tryptophan (cyclo-GW) [120], cyclo-phenylalanine tryptophan (cyclo-FW) [121], and bicyclic β-peptides [122]. Among these, FF, with a piezoelectric coefficient d₃₃ of 18 ± 5 pm/V, is one of the most extensively researched piezoelectric peptides owing to its straightforward structure, robust mechanical properties, and pronounced piezoelectric properties [122]. Structurally, FF comprises two phenylalanines molecules (Phe) linked as NH₂-Phe-Phe-COOH [123].

A defining characteristic of peptides is their self-assembly capability into ordered nanostructures through intermolecular interactions, including hydrogen bonding, electrostatic forces, solvent-mediated forces, and π-π stacking. These interactions yield nanostructures such as nanowires and nanotubes that exhibit excellent thermal and mechanical stability, semiconductor properties, piezoelectricity, and optical characteristics [[122], [123], [124]].

3.2.7. Collagen

Collagen, a protein widely recognized for furnishing mechanical structure support in biological systems [125], also exhibits piezoelectric properties, with coefficient typically ranging from 0.2 to 2.0 pC/N [126]. Its exceptional biocompatibility, excellent cell adhesion, hydrophilicity, low antigenicity, and biodegradability make collagen an ideal candidate for use as an intelligent piezoelectric biomaterial, particularly for cartilage regeneration [127]. For instance, calcium phosphate-containing collagen (CaP-Gelfix) composite scaffolds have exhibited remarkable potential in promoting the regeneration and repair of cartilage tissue. Studies have demonstrated that this scaffold achieved an average fill rate of newly formed cartilage tissue in the defective area of approximately 81 % and 96 % at 8 and 20 weeks, respectively, with minimal foreign body reaction, thereby effectively promoting cartilage regeneration [96]. Collagen's relatively low d₃₃ value stems from its primarily axial, rather than radial, piezoelectric response [51]. Other limitations of collagen include low mechanical strength, rapid degradation rates, and the use of potentially toxic cross-linking agents during their processing [128].

3.3. Piezoelectric composites

Piezoelectric composites are typically fabricated by embedding piezoelectric particles or fibers into a matrix material, which can be composed of polymers, metals, or ceramics. This section focuses on smart piezoelectric composites that can be degraded in living organisms, such as piezoelectric hydrogels (Table 1). These materials have merged as a preferred choice for cartilage regeneration scaffolds due to their ability to overcome the brittleness associated with inorganic materials while retaining high piezoelectric performance and incorporating the advantages of organic materials. For example, Vinikoor et al. developed an ultrasound-responsive, biodegradable, and injectable piezoelectric hydrogel comprising cryosectioned PLLA piezoelectric short nanofibers embedded in a collagen matrix. This hydrogel can be injected into the body, adapt to various shapes to repair severe cartilage defects, and fully degrade once tissue repair was achieved [74]. This illustrates the significant potential of hydrogel-based piezoelectric materials for non-invasive cartilage repair [129,130].

Compared to pure organic or inorganic piezoelectric materials, piezoelectric composites offer the advantage of tunable piezoelectric performance, which can be adjusted through various processing technique. Additionally, they offer benefits such as adjustable structural gradients, ease of biodegradation, injectability, and high biocompatibility[131].

Intelligent piezoelectric composites can also exhibit multifunctional properties, such as antimicrobial activity, ROS-releasing or scavenging properties, immunomodulatory effects, and the ability to modulate the vascular microenvironment [132]. For example, wu et al. engineered a bilayer-structured piezoelectric hydrogel to mimic the tissue architecture of osteochondral bone. They discovered that the piezoelectric effect of this biomimetic hydrogel not only promoted cartilage regeneration but also exhibited antimicrobial properties [133]. Furthermore, intelligent piezoelectric composites have been shown to induce macrophage polarization from pro-inflammatory M1 phenotype to the anti-inflammatory M2 phenotype [75]. Transcriptome sequencing analysis revealed that this immunomodulatory effect is correlated with the phosphatidylinositol-3-kinase (PI3K) / protein kinase B (Akt) signaling axis [70,75]. These composites also exhibit significant potential for recruiting stem cells to promote tissue repair and facilitate cartilage healing [134]. For instance, Ricotti et al. developed a composite hydrogel containing piezoelectric BaTiO₃ nanoparticles and embedded adipose-tissue-derived mesenchymal stromal cells. Upon US stimulation, the composite hydrogel efficiently initiated cartilage regeneration in degraded cartilage tissue [24].

Another notable advantage of intelligent piezoelectric composites lies in their ability to spatially edit intrinsic structure. Ding et al. designed functionalized void hyaluronic acid microcarriers (HAMCs) using an emulsion process combined with the vacuole method. These microcarriers significantly enhanced the secretion of chondrocyte proteoglycans and type II collagen compared to conventional HAMCs and 2D planar materials. Moreover, they inhibited the Wnt/catenin signaling pathway during long-term culture, thereby preventing the dedifferentiation of transplanted chondrocytes and maintaining the stability of the chondrocyte phenotype [135].

Given cartilage's heightened sensitivity to ES relative to other tissues, piezoelectric hydrogels are particularly well-suited for cartilage tissue engineering and regenerative medicine applications [136,137]. These flexible materials boast injectable, light-curing, biocompatible, and degradable properties [138]. Despite their numerous advantages, challenges such as swelling behavior, rapid degradation, and structural deformation must be addressed in future applications [139].

4. Advances in piezoelectric biomaterials for cartilage regeneration

4.1. Advances in the forms of piezoelectric biomaterials

Piezoelectric materials used for cartilage regeneration are available in a variety of structural forms, including 1D nanoparticles, flexible 2D films and microsphere hydrogels, 3D scaffolds and bioink, 4D polymers with temperature responsiveness and shape memory (Fig. 4). Cai et al. prepared hollow ZnO (HZnO) nanohollow spheres using a one-step method to facilitate the enhancement of the piezoelectric effect and to achieve tumour-targeting capability of HZnO by ultrasonic mechanical forces [140]. A key consideration in the use of inorganic nanoparticles is the potential risk of nanoparticle accumulation, which may pose a great health risk [141]. Using different weaving methods, Kim et al. found that optimised pressure sensors with a 2/2 weft rib pattern produced a high sensitivity of 83 mV/N, 245 % higher than that of the 1/1 pattern, for the same 2D film PVDF [142]. 2D piezoelectric biofilms offer flexibility, ease of production, and cost-effectiveness, making them an ideal alternative to biological films such as fascia and perichondorm [143]. However, studies have shown that 2D planar materials exhibit limited efficacy in promoting cartilage regeneration. In contrast, 3D porous structures offer significant spatial advantages. Their structures can be realized through bioprinting technique, which combines bioinks containing living cells or cytokines with traditional biomaterials. A major challenge of this approach is to avoid inactivation of living cells and cytokines during the printing process, as this can significantly compromise the functionality and efficacy of the final constructs [[144], [145], [146]]. Similarly, Ghosal et al. developed a 4D-printed polymer with temperature-responsive and shape-memory properties. This material also exhibits excellent antimicrobial properties against both gram-negative and gram-positive bacteria, as well as mechanical properties comparable to those of ligaments and articular cartilage [147].

Fig. 4.

Advances in piezoelectric materials from 1D to 4D. (A) Nanoparticle dimensional synthesis of piezoelectricity-enhanced degradable and hollow ZnO heterostructures with the possibility of generating more ROS (reproduced with permission from ref.[140] Copyright 2013 ACS); (B) preparation of piezoelectric PVDF films by 2D electrostatic spinning technology (reproduced with permission from ref.[142] Copyright 2022 NPJ Flex. Electron.); (C) bioink was prepared by mixing C28/I2 chondrocytes with GelMA/SG solution, printed to the corresponding shape by a 3D bioprinter for in vitro culture experiments, and then implanted subcutaneously in nude mice for in vivo chondrogenesis experiments (reproduced with permission from ref.[144] Copyright 2024 Elsevier); (D) 4D polymers with temperature responsiveness and shape memory prepared by 3D printing technology (reproduced with permission from ref.[147] Copyright 2024 WILEY).

4.2. Advances in the multifunctional piezoelectric composites

Multifunctional piezoelectric biomaterials can be engineered to incorporate drugs, growth factors and signaling molecules, or stem cells (Fig. 5). These bioactive components can be released in a controlled manner, such as through gradual release during material degradation, to promote cell adhesion, stem cell differentiation, and angiogenesis, thereby promoting tissue regeneration and growth [148,149]. For example, Yue et al. used piezoelectric materials in conjunction with acoustic kinetic therapy for the treatment of tumors [150].

Fig. 5.

Application prospect of piezoelectric biomaterials in cartilage regeneration engineering. (A) Ultrasound-activated drug release (reproduced with permission from ref.[149] Copyright 2024 KeAi); (B) combination of sonodynamic therapy and piezoelectric materials (reproduced with permission from ref.[150] Copyright 2024 WILEY); (C) biodegradable piezoelectric conductive integrated hydrogel scaffolds are used to repair cartilage defects under mechanical vibration (reproduced with permission from ref.[152] Copyright 2024 WILEY); (D) scheme of 3D bioprinting process of cartilage engineering (thermally induced phase separation) (Reproduced with permission from ref.[146] Copyright 2020 Elsevier); (E) magnetic microcarriers for cartilage defect repair (reproduced with permission from ref.[153] Copyright 2023 ACS); (F) processes of Cartilage-on-a-chip assembly and cell perfusion (reproduced with permission from ref.[71] Copyright 2024 KeAi).

In addition, bioelectricity generated by piezoelectric materials can directly apply polarized charges to substances within the tissue microenvironment, facilitating carrier separation and inhibition of complexation process [151,152]. Ma et al. developed magnetic microcarriers for cartilage defect repair by in-situ polymerization of dopamine (DA) and Fe₃O₄ to form a complex. These microcarriers not only exhibit porosity suitable for cell growth, excellent hemocompatibility and biocompatibility, but also possess magnetic responsiveness, enabling deformation under an applied magnetic field [153].

In pharmaceutical engineering, cell perfusion technology can also be integated with cartilage microarrays to create regenerative cartilage models. These models simulate the effects of various stimuli on cartilage tissue in vivo, providing a valuable platform for studying cartilage regeneration and drug responses [71].

4.3. Advances in artificial intelligence

Artificial intelligence (AI) can visualize ultrasonic propagation in piezoelectric materials. Coulomb-coupled imaging methods are inefficient in terms of time-intensity due to the time-consuming nature of conventional point-by-point scanning. Banerjee et al. used Generative Adversarial Networks (GANs) for data augmentation, achieving AI-driven predictive modelling and real-time adaptive imaging [154]. Whereas X-ray computed tomography typically takes 5 h for a volume of 100 cubic millimetres, ultrasound imaging allows real-time imaging at high speeds of 25–100 fps. However, ultrasonic sensors tend to have poor sensitivity. Therefore, Karthikeyan et al. used an explainable AI (XAI) method called local interpretable model-agnostic explanations (LIME) to improve the accuracy to more than 80 % [155]. In the strategy of US stimulation of piezoelectric materials for cartilage regeneration, AI also has great potential for application in guiding the synthesis of piezoelectric materials. Yang et al. used a deep learning model as an inversion tool for resonance ultrasound spectroscopy of piezoelectric materials, which can efficiently determine the complete tensor constants of the materials [156]. Hu et al. trained a predictive model of piezoelectric coefficients using machine learning and deep graph neural network-based auto-learning, predicted the piezoelectric coefficients of 12,680 materials, and reported on the top 20 potential high-performance piezoelectric materials [157].

AI can also predict the electric field distribution in piezoelectric materials and cartilage tissue under ultrasound stimulation. Ricotti et al. proposed an analytical model that predicts the voltage generated by piezoelectric nanoparticles controlled by ultrasound. By aggregation of nanoparticles in chondrocytes, the model can also be used to predict the distribution of electric field flow lines in the cytoplasm [24]. Mathiessen et al. developed AI for assessing cartilage thickness and joint inflammation [158]. This AI-guided quantitative ultrasound assessment could become the new standard for monitoring cartilage health, facilitating early detection of cartilage lesions and enabling timely intervention. Fiorentino et al. also developed a deep learning framework for assessing cartilage thickness by characterizing two interfaces of cartilage and calculating cartilage thickness [159].

5. Challenge and perspective

In summary, this review comprehensively reviews the integration of ultrasound with piezoelectric biomaterials for applications in cartilage tissue repair. A deep understanding of the evolution of cellular properties and their regulatory mechanisms during cartilage tissue regeneration, particularly under ultrasound-activated piezoelectric stimulation, is essential for advancing therapeutic strategies. While the potential of this approach has been widely recognized, several challenges and opportunities remain for tis practical implementation. Key points include:

i) Optimization of US parameters and piezoelectric material properties: The frequency, intensity and duration of ultrasound significantly influence the activation efficiency of piezoelectric materials. Optimizing these therapeutic parameters, along with the properties of the piezoelectric biomaterials, will facilitate the development of advanced models for cartilage tissues under ultrasound stimulation, such as cartilage organoids. Although direct integration of cartilage chips with biocomputers has not yet been achieved, the theoretical development of computing systems based on chondrocytes or similar tissues represents a promising future direction [160].

ii) Regulation of cartilage regeneration: In-depth studies on the temporal regulation of cartilage regeneration are essential to ensure adequate repair in the absence of blood supply, preventing issues such as over-proliferation, dedifferentiation, or fibrosis. Such research will promote the development of ultrasound-stimulated piezoelectric biomaterials for regenerative medicine applications.

iii) Multimodal synergistic therapeutic strategies: Ultrasound-activated piezoelectric materials can be combined with other therapeutic approaches, such as drug delivery, gene therapy, cell therapy, sonodynamic therapy (SDT) and 3D printing, to create multimodal synergistic strategies. For example, ultrasound can act as a triggering source for drug delivery systems, working in tandem with piezoelectric materials to enhance localized drug concentrations.

iv) Stem cell integration for enhanced cartilage repair: Stem cells possess the unique ability to self-renew and differentiate into specific cell types, which can promote cartilage repair. Ultrasound-activated piezoelectric biomaterials can further accelerate cartilage regeneration by modulating the behavior of stem cells through mechanical and electrical stimulation.

v) Artificial intelligence (AI) in piezoelectric material design: AI technology allows the development of materials possessing high piezoelectricity and high biocompatibility. While traditional material design often relies on trial and error, AI can accelerate the development of novel piezoelectric materials by using machine learning algorithms to simulate and predict the material properties [161]. In addition, AI also has great potential for assessing cartilage tissue damage and regeneration in ultrasound piezoelectric therapy cartilage regeneration strategies.

In conclusion, a deeper understanding of the complex biological interactions underlying ultrasound-stimulated piezoelectric biomaterials will enhance the precision of treatment protocols. Advances in AI, organ-on-a-chip, stem cells, and piezoelectric biomaterial design will pave the way for more precise and personalized approaches to cartilage regeneration and repair. These innovations will accelerate clinical translation and ultimately improve treatment outcomes for patients.

CRediT authorship contribution statement

Yangchen Wei: Writing – review & editing, Writing – original draft, Formal analysis, Data curation, Conceptualization. Zhengyang Li: Writing – original draft, Software, Project administration, Methodology, Investigation. Tianjing Yu: Software, Resources, Formal analysis, Data curation. Yan Chen: Software, Resources, Formal analysis, Data curation. Qinglai Yang: Visualization, Supervision, Methodology. Kaikai Wen: Writing – review & editing, Supervision, Funding acquisition. Junlin Liao: Visualization, Validation, Supervision, Funding acquisition. Linlin Li: Writing – review & editing, Supervision, Funding acquisition.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.